Microneedle arrays for biosensing and drug delivery

ABSTRACT

Methods, structures, and systems are disclosed for biosensing and drug delivery techniques. In one aspect, a device for detecting an analyte and/or releasing a biochemical into a biological fluid can include an array of hollowed needles, in which each needle includes a protruded needle structure including an exterior wall forming a hollow interior and an opening at a terminal end of the protruded needle structure that exposes the hollow interior, and a probe inside the exterior wall to interact with one or more chemical or biological substances that come in contact with the probe via the opening to produce a probe sensing signal, and an array of wires that are coupled to probes of the array of hollowed needles, respectively, each wire being electrically conductive to transmit the probe sensing signal produced by a respective probe.

CROSS REFERENCE TO RELATED APPLICATIONS

This patent application is a continuation of, and claims priority andbenefits of, U.S. patent application Ser. No. 14/342,536 filed Jul. 30,2014, which is a 35 USC §371 National Stage application of InternationalApplication No. PCT/US2012/053544 filed Aug. 31, 2012, which furtherclaims the benefit of priority of U.S. Provisional Application No.61/530,927, filed on Sep. 2, 2011. The entire content of thebefore-mentioned patent applications is incorporated by reference aspart of the disclosure of this application.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

This invention was made with government support under grants no.N00014-08-1-1202 awarded by the Office of Naval Research (ONR), grantno. DE-AC04-94AL85000 awarded by the U.S. Department of Energy to SandiaCorporation, and grant no. 151337 awarded by Sandia NationalLaboratories Laboratory Directed Research and Development (LDRD). Thegovernment has certain rights in the invention.

TECHNICAL FIELD

This patent document relates to biosensors and drug delivery devices.

BACKGROUND

Sensing biological events in vitro and in vivo can provide real-timedetection of physiologically relevant compounds, such as monitoring ofmetabolites, electrolytes, biochemicals, neurotransmitters, medicallyrelevant molecules, cancer biomarkers, and pathogenic microorganisms.Devices that perform such biological event sensing are known asbiosensors, which can provide real-time detection of physiologicalsubstances and processes in living things. A biosensor is an analyticaltool that can detect a chemical, substance, or organism using abiologically sensitive component coupled with a transducing element toconvert a detection event into a signal for processing and/or display.Biosensors can use biological materials as the biologically sensitivecomponent, e.g., such as biomolecules including enzymes, antibodies,nucleic acids, etc., as well as living cells. For example, molecularbiosensors can be configured to use specific chemical properties ormolecular recognition mechanisms to identify target agents. Examples caninclude evaluating physiologic and pathologic activity within a tissue,as well as drug discovery and drug screening. Biosensors can use thetransducer element to transform a signal resulting from the detection ofan analyte by the biologically sensitive component into a differentsignal that can be addressed by optical, electronic or other means. Forexample, the transduction mechanisms can include physicochemical,electrochemical, optical, piezoelectric, as well as other transductionmeans.

SUMMARY

Techniques, systems, and devices are disclosed for biosensing andtherapeutic interventions.

In one aspect of the disclosed technology, a device includes an array ofhollowed needles, in which each needle includes a protruded needlestructure including an exterior wall forming a hollow interior and anopening at a terminal end of the protruded needle structure exposing thehollow interior, and a probe inside the exterior wall to interact withone or more chemical or biological substances that come in contact withthe probe via the opening to produce a probe sensing signal, and anarray of wires that are coupled to the probes of the array of hollowedneedles, respectively, each wire being electrically conductive totransmit the probe sensing signal produced by a respective probe.

Implementations can optionally include one or more of the followingfeatures. For example, the one or more of the probes can include afunctionalized coating configured to interact with an analyte within afluid. An electrochemical interaction between the analyte and thecoating on one of the one or more functionalized probes can be detectedusing at least one of amperometry, voltammetry, or potentiometry. Thedevice can further include a processing unit in communication with thearray of wires that receives the probe sensing signals and uses theprobe sensing signals as data. The processing unit can compare the datato a threshold value to determine whether the analyte concentrationreflects a healthy or disease state. The processing unit can determine apattern in the data that indicates whether the analyte concentrationreflects a healthy or disease state. The processing unit can multiplexthe received probe sensing signals from the probes. The device can beintegrated into an adhesive patch for placement on skin to detect theanalyte residing in transdermal fluid.

In another aspect of the disclosed technology, a device includes asubstrate that includes a microneedle with a hollowed interior locatedon one side of the substrate, in which the microneedle includes a wallwith an opening to the hollowed interior, an electrode including aprobe, in which the probe is disposed inside the hollowed interior, anda wire that is connected to the probe, in which the electrode isfunctionalized by a coating over the probe to interact with an analyteto produce an electrical signal.

Implementations can optionally include one or more of the followingfeatures. For example, an electrochemical interaction between theanalyte and the coating on the functionalized electrode can be detectedusing at least one of amperometry, voltammetry, or potentiometry. Thedevice can further include a processing unit in communication with thewire that receives the electrical signal and uses the electrical signalas data. The processing unit can compare the data to a threshold valueto determines whether the analyte concentration reflects a healthy ordisease state. The processing unit can determine a pattern in the datathat indicates whether the analyte concentration reflects a healthy ordisease state. The device can be integrated into an adhesive patch forplacement on skin to detect the analyte residing in transdermal fluid.The device can further can include a polymer film having pores of areversibly tunable porosity, in which the polymer film is attached to anopposite side of the substrate, a protrusion structure configured on theone side of the substrate, in which the protrusion structure has achannel between an opening in the substrate exposing the polymer filmand an opening at a terminal end of the protrusion structure, acontainment structure that contains a chemical substance, in which thecontainment structure includes one or more openings attached to thepolymer film positioned above the protrusion structure, and an electrodeattached to the polymer film, in which the electrode provides anelectrical stimulus to trigger an expansion of the pores of the polymerfilm to an open state or a contraction the pores of the polymer film toa closed state. The processing unit can be in communication with thewire that receives the electrical signal to use as data and incommunication with the electrode to generate the electrical stimulus.The processing unit can process the data to determine whether theanalyte concentration reflects a healthy or disease state. Theprocessing unit can actuate the electrode to apply an electricalstimulus to the polymer film to alter its permeability from the closedstate to the open state, thereby releasing the chemical substance fromthe device. The processing unit can multiplex the received electricalsignals and the actuation of the electrical stimuli.

In another aspect, a method to sense an analyte and deliver atherapeutic agent includes detecting a signal produced by an analyte atan interface with a chemically functionalized probe configured toelectrochemically interact with the analyte within a biological fluid,in which the signal is transduced to an electrical signal by thechemically functionalized probe, processing the electrical signal todetermine a parameter of the analyte, and based on the determinedparameter, applying an electrical stimulus to a valve comprising aporous polymer film having pores of a reversibly tunable porosity, thevalve attached to a container containing a therapeutic agent, in whichthe electrical stimulus alters the permeability of the pores from aclosed state to an open state, thereby releasing the therapeutic agentinto the biological fluid.

In another aspect, a device includes a substrate that includes aplurality of microneedles with a hollowed interior located on one sideof the substrate, in which each of the microneedles includes a wall withan opening to the hollowed interior, a biosensor module, an actuatormodule, and a processing unit in communication with the plurality ofwires to receive the electrical signal and use the received electricalsignal as data, in which the processing unit is in communication withthe actuator electrode to generate the electrical stimulus based on thedata. The biosensor module includes a plurality of sensing electrodesdisposed inside the hollowed interior of a first group of the pluralityof microneedles, the sensing electrodes including a probe, in which theprobe includes a functionalized coating configured to interact with ananalyte within a fluid to produce an electrical signal, and a pluralityof wires, in which one wire of the plurality of wires is connected tothe probe of the sensing electrodes. The actuator module includes apolymer film having pores of a reversibly tunable porosity, in which thepolymer film is attached to an opposite side of the substrate, aplurality of protrusion structures disposed inside the hollowed interiorof a second group of the plurality of microneedles, in which theprotrusion structures includes a channel between an opening in thesubstrate exposing the polymer film and an opening at a terminal end ofthe protrusion structure, a containment structure that contains achemical substance positioned above the polymer film, in which thecontainment structure includes one or more openings coupled to thepolymer film positioned above the protrusion structure, and an actuatorelectrode attached to the polymer film, in which the actuator electrodeprovides an electrical stimulus to trigger an expansion of the pores ofthe polymer film to an open state or a contraction the pores of thepolymer film to a closed state.

Implementations can optionally include one or more of the followingfeatures. For example, the processing unit can compare the data to athreshold value to determines whether the analyte concentration reflectsa healthy or disease state. The processing unit can determine a patternin the data that indicates whether the analyte concentration reflects ahealthy or disease state. The processing unit can actuate the actuatorelectrode to apply the electrical stimulus to the polymer film to alterits permeability from the closed state to the open state, therebyreleasing the chemical substance into the fluid. The processing unit canmultiplex the received electrical signals from the probe and theactuation of the electrical stimuli to the actuator electrode. Theprocessing unit can include logic gates configured on the substrate. Thedevice can be integrated into an adhesive patch for placement on skin todetect the analyte residing in transdermal fluid.

The subject matter described in this patent document can be implementedin specific ways that provide one or more of the following features.Microneedle array devices and techniques are described for performingmultiplexed sensing applications and/or drug delivery in an autonomous,minimally-invasive, and controlled manner. For example, the disclosedtechnology can be implemented to detect analytes in living things viaelectrochemical methods using microneedle arrays that can be integratedinto a patch and applied to the skin. Biosensing can be implementeddirectly at the microneedle-transdermal interface without the uptake andsubsequent processing of biological fluids. Potentiometric,voltammetric, and amperometric techniques can be used to transducephysiological and biochemical information using the microneedle arrayplatform, which can be integrated into one, all-inclusive platform toenable direct biosensing of multiple analytes in bodily fluids.Additionally, the biosensing functionality can be coupled with actuationfunctionality. For example, a therapeutic agent (e.g., drugs, vaccines,insulin, hormones, vitamins, anti-oxidants, and other pharmacologicalagents) delivery feature can be initiated by stimuli-responsiveconducting polymer nanoactuators. The biosensor-actuator platform can beintegrated on an adhesive patch to monitor key physiological/biochemicalparameters and/or deliver a therapeutic intervention on demand. Theadhesive patch can be integrated with electronics to allow signaltransduction and communication. The technology can be used as a “sense”constituent and as a “treat” constituent in an exemplary“Sense-Act-Treat” feedback loop process, which can be utilized in avariety of applications that can include, at least, wireless healthcare,personalized medicine, health profiling, performance/health monitoring,and athlete/warfighter monitoring.

For example, the disclosed technology can have wide-ranging applicationswithin a multitude of fields and disciplines where the assessment ofhealth in real-time is desired. For example, the technology can beeasily applied for use in the generalized healthcare, fitness, sport,remote monitoring, wireless healthcare, personalized medicine,performance/health monitoring, and warfighter monitoring domains. Theminimally-invasive nature of the technology, combined with its robustarchitecture, can make the technology well-suited for diverse biomedicalmonitoring applications, e.g., obtaining biomarker signatures for healthprofiling, or patterns of bioanalytes as a measure ofperformance/fitness. As another example, cancer cells, such as melanoma,are known to undergo increased levels of glycolysis which causelocalized environments of decreased pH and glucose concentrations, andincreased lactate concentrations; thus the disclosed technology cansimultaneously and locally detect glucose, lactate, and pH, thereby canbe used as a point-of-care clinical diagnostic device to determine ifskin cells are cancerous and give immediate data before a lengthy biopsycan be performed. For example, when the technology is used as the“sense” constituent in the exemplary ‘Sense-Act-Treat’ feedback loop,the technology can be employed as an element of a smart patch that isable to trend pertinent physiological/biochemical information forhigh-risk patients (e.g., stroke, cardiac, etc.). Moreover, for example,this feature of the technology can be adapted as a “battlefieldhospital-on-a-patch” that is able to determine the occurrence of acuteinjury/trauma and alert the appropriate personnel to instigate a rapidevacuation of the individual and begin a targeted treatment regimen.When the technology is used as the “treat” constituent in the exemplary‘Sense-Act-Treat’ feedback loop, the technology can be employed as anelement of the smart patch that is able to provide a targeted therapyfor acute events experienced by high-risk patients (stroke, cardiac,etc.). These exemplary features can be adapted as a “battlefieldhospital-on-a-patch” that is able to begin a treatment regimen in combatsituations where the rapid evacuation and treatment of injured personnelis not feasible.

BRIEF DESCRIPTION OF THE DRAWINGS

FIGS. 1A-1C show schematics of an exemplary microneedle array device.

FIGS. 1D-1F show images of an exemplary microneedle and microneedlearray.

FIGS. 1G-1I show images of an exemplary microneedle array device on anadhesive patch implemented on a living being.

FIG. 1J shows an exemplary diagram of a ‘Sense-Act-Treat’ feedback loop.

FIG. 1K shows a diagram of an exemplary microneedle arraysensor-actuator featuring individually-addressable microneedles tomonitor for different dysfunctions and individually-addressablemicroneedles to deliver a therapeutic agent in response to detection ofthe dysfunctions.

FIG. 1L shows an illustration of an exemplary multiplexed controlledrelease of targeted therapeutic agents.

FIGS. 2A-2C show schematics of an exemplary microneedle array stripsystem.

FIGS. 3A-3C show data plots of exemplary results of multiplexeddetection of various biomarkers on an electrochemical array strip usingindividual microneedles.

FIGS. 4A and 4B show exemplary microneedle arrays fabricated usingscreen printing and stencil processes to create a pattern.

FIG. 5A shows an illustration of an exemplary process to fabricate anexemplary bicomponent microneedle electrode array using solid and hollowmicroneedle constituents.

FIG. 5B shows a schematic of the exemplary bicomponent microneedleelectrode array fully-assembled.

FIG. 5C shows an illustration of an exemplary process to grow aglutamate oxidase (GluOx)-functionalized poly(o-phenylenediamine) (PPD)film at the solid microneedle surface.

FIG. 5D shows an illustration of the biocatalytic behavior of theexemplary electropolymerized glutamate oxidase-poly(o-phenylenediamine)film.

FIGS. 6A and 6B show scanning electron microscopy (SEM) images ofexemplary solid and hollow microneedle arrays, respectively.

FIG. 7 shows a hydrodynamic voltammogram data plot of exemplaryglutamate bicomponent microneedle array electrodes.

FIGS. 8A and 8B show exemplary chronoamperogram data plots recorded forincreasing concentrations of glutamate.

FIG. 9 shows exemplary chronoamperogram data plots recorded in 0.1 Mphosphate buffer.

FIG. 10 shows exemplary data plots showing the stability of theglutamate response over extended time periods.

FIGS. 11A and 11B show data plots showing sensitivity of an exemplaryglucose microneedle biosensor.

FIGS. 12A and 12B show a schematic illustration of an exemplarymicroneedle-based multi-channel, multiplexed drug delivery actuatordevice.

FIGS. 13A and 13B show SEM images of the surface morphology of anexemplary hollow microneedle array.

FIG. 14 shows images of triggered release of methylene green from theindividually-addressable reservoirs of an exemplary microneedle-baseddrug delivery actuator device.

FIG. 15 shows time-lapse still frame images of the release of methylenegreen dye from a single microneedle of an exemplary microneedle-baseddrug delivery actuator device.

FIG. 16 shows an exemplary UV-Vis spectrum data plot illustrating theabsorbance for the release of methylene green dye from an exemplarymicroneedle.

FIG. 17 shows a schematic of an exemplary microneedle during drugdelivery.

FIGS. 18A-18D show illustrative schematics showing processing steps forthe assembly of an exemplary microneedle array device.

FIGS. 19A and 19B show optical images of an array of carbon fiberelectrodes and a single carbon fiber electrode.

FIGS. 20A and 20B show SEM images of an exemplary hollow microneedlearray.

FIG. 21A shows an image of skin after application of an exemplarymicroneedle array.

FIG. 21B and FIG. 21C show optical micrographs of hollow microneedlesbefore and after insertion into skin.

FIGS. 22A and 22B show SEM images of an exemplary hollow microneedlearray.

FIG. 23 shows a data plot of a cyclic voltammetric scan of ferricyanidein KCl versus Ag/AgCl reference and Pt counter electrodes.

FIG. 24 shows a data plot of cyclic voltammetric scans of hydrogenperoxide versus Ag/AgCl reference and Pt counter electrodes.

FIG. 25 shows a data plot of voltammetric scans of buffer solution andascorbic acid in buffer versus Ag/AgCl reference and Pt counterelectrodes.

FIGS. 26A and 26B show optical micrographs of the unpacked and Rh-carbonpaste packed microneedle array.

FIGS. 27A and 27B show SEM images of the unpacked and Rh-carbon pastepacked microneedle constituent of the array.

FIG. 28A shows plots of hydrodynamic voltammograms of buffer andhydrogen peroxide at the rhodium-dispersed carbon paste microneedleelectrode.

FIG. 28B shows plots of chronoamperograms obtained using the exemplaryrhodium-dispersed carbon paste microneedle electrode.

FIG. 29 shows a plot of a calibration curve obtained for hydrogenperoxide concentrations.

FIG. 30A shows plots of chronoamperograms obtained for lactate.

FIG. 30B shows a plot of a calibration curve obtained for lactateconcentrations.

FIG. 31 shows plots of chronoamperograms showing the effect ofphysiologically-relevant electroactive interferents.

FIG. 32 shows a data plot showing the stability of the electrochemicalresponse of an exemplary microneedle array used for lactate detection.

DETAILED DESCRIPTION

Techniques, systems, and devices are disclosed for the detection ofanalytes and delivery of therapeutic compounds in living things usingmicroneedle array based biosensors and actuators.

In one aspect, the present technology includes a device using an arrayof microscale structures that penetrate the surface of a biologicaltissue to detect fluctuations in certain biomarkers in tissue fluidsand/or extracellular fluids. By detecting such fluctuations, the devicescan be used to monitor progression of diseases, illnesses, and acuteinjuries, among other conditions. For example, this can be implementedby loading the microstructures with electrochemical transducers in theform of microscale needles, e.g., also referred to as microneedles,microprobes, electrodes, or probes, which can have different chemicalfunctionalities towards biochemical and physiological analytes, e.g.,such as a biochemical, metabolite, electrolyte, ion, pathogen,microorganism, etc. For example, the device can employ variouselectrochemical techniques to perform electrochemical reactions directlyat the microneedle/fluid interface and transduce that biochemicalinformation into an electrical signal (e.g., voltammetric,potentiometric, amperometric, conductometric, and/or impedimetric),which can be further processed. For example, each microneedle in anarray can be configured to detect a different analyte and multiplexed tobe addressed by one or more transduction modalities, e.g., such as anN-array of microneedle elements, in which each element N1 sensesglucose, N2 senses lactate, element N3 senses fatty acids, etc., andeach analyte is transduced using any or all of voltammetric,potentiometric, amperometric, conductometric, and/or impedimetrictechniques.

The device can also be used to implement a therapeutic intervention thatutilizes the microneedles to release a chemical agent (e.g., a drug)into the fluid in a controlled manner at a particular localized area inwhich the microneedle has been applied. The delivery of a targetedtherapeutic intervention can be implemented in response to an acuteevent or based on a chronic condition, e.g., monitored by the sensingcontingent of the device. The microneedle array can be configured inconjunction with a permeability-tunable conducting polymer material tocontrol the porosity of the polymer. For example, the device can beconfigured to include one or more reservoirs containing the chemicalagents(s) coupled to the permeability-tunable conducting polymermaterial that is positioned between the reservoir(s) and the substrateof the microneedle array. Under certain electrochemical stimuli, thepolymer material can selectively be made porous (e.g., change theporosity of the polymer material), which can effectively act as a valvethat can be selectively opened or closed to transport the chemical agentfrom the reservoir through the microneedle (lumen) and subsequently intothe tissue fluid. The chemical agent release mechanism iselectrochemically enabled, e.g., without moveable parts ormicroelectromechanical (MEMS) components.

The actuation of the therapeutic contingent of the exemplary device canbe controlled using an integrated logic system or processing unit, whichcan provide an electrical stimulus based on feedback from the sensingcontingent of the exemplary device. For example, the described sensingcontingent of the exemplary device can continuously monitor aconcentration level (e.g., based on fluctuations in an average level, amaximum or minimum threshold level, etc.) of a particular analyteassociated with normal function or dysfunction indicative of diseases,illnesses, and acute injuries or other conditions and transduce thedetected biochemical information associated with the analyte into anelectrical signal. The electrical signal can be processed using aprocessing unit, which can include, at least, a processor and a memorycoupled to the processor. For example, the memory may encode one or moreprograms that cause the processor to perform one or more of the methodacts described in this patent document, e.g., including storing thedetected signals, analyzing the detected and/or stored signals againstother stored values (e.g., such as analyte threshold values indicativeof a healthy or dysfunctional state) and/or determining whether or notto release a chemical agent using the therapeutic contingent of theexemplary device. For example, the processing unit may determine thedetected analyte level has exceeded a threshold value stored in thememory, and subsequently activate the described actuator to release adrug in response to the determined analyte level. For example, this canbe performed by applying a suitable redox potential, the exemplarydevice can “open” and “close” the described polymer in a reversiblemanner by changing the intrinsic porosity of the matrix, thus triggeringthe flux of medication from an on-body reservoir directly into thetransdermal fluid. Multivariate/multiplexed drug delivery can be used toimplement a therapy in a unique manner, where drugs can be delivered ateach microneedle constituent of the array. For example, owing to thearrayed nature of the microneedle structures, multivariate/multiplexeddrug delivery can be realized, and a unique analyte can be detected ateach microneedle constituent of the array or multiple analytes can bedetected at each individual needle through an array of electrodes withina needle. Moreover, the arrayed nature of an exemplary system can enablethe device to tailor the cocktail of drugs to mitigate various forms ofinjury/trauma. Furthermore, the ability to selectively control theporosity of the membrane by adjusting the applied redox potential canimply that the flux rate, and hence the dosage, can be controlled asneeded via the integrated biosensor and/or the logic-gate sensing andprocessing unit.

In some examples, the described microneedle biosensor-actuatortechnology can be implemented transdermally by applying an exemplarymicroneedle device to the skin. In other examples, the describedmicroneedle biosensor-actuator technology can be implemented in vivo toother organs within the body, e.g., including the liver, the sclera ofthe eye, etc. In an example of the biosensing functionality of thedisclosed technology, a device can include an array of microneedlessensor-actuators integrated on a skin adhesive patch and applied to theskin of a living being to transdermally monitor physiological andbiochemical parameters (e.g., glucose). In an example of blood glucosemonitoring, the microneedles can be functionalized with glucose oxidaseenzyme (a biocatalyst) that is entrapped within a conducting polymer,e.g., in which the electrode component is conductive and functionalized(e.g., coated) to include the biocatalyst. Upon application of the patchto the skin, the microneedles penetrate the skin so that extracellularfluid (e.g., blood) can diffuse into the microneedle. The biocatalyst,as glucose diffuses into, can convert the glucose substrate intogluconic acid. In the meantime, since the conversion of glucose intogluconic acid is an oxidation-reduction (redox) reaction (e.g.,oxidizing glucose), reduction also occurs to oxygen and water naturallypresent in the blood to form hydrogen peroxide. Hydrogen peroxide is anelectrochemically active species, which can be oxidized or reduced atcertain potentials at the electrode. For example, this can be doneamperometrically, in which a potential is applied and current ismonitored, or voltammetrically, in which the potential is changed andcurrent change is monitored. For example, as the hydrogen peroxidechanges at the electrode in which a potential is applied, thecorresponding current change is monitored as an electrical signal thatcan be further processed using signal processing techniques.

In one embodiment, a minimally-invasive multi-component microneedledevice for detecting an analyte and delivering a therapeutic compoundcan include a microneedle array in conjunction with electrodes, e.g.,which can be chemically-functionalized, enzyme-functionalized, and/orion-selective electrodes, to perform multiplexed sensing and actuatingapplications in an autonomous, minimally-invasive, and controlledmanner.

FIGS. 1A and 1B show schematics of an exemplary device based on hollowedneedles with probes. FIG. 1A shows a schematic of an exemplarymicroneedle array device 100, and FIG. 1B shows an exemplary schematicof the device unassembled. The microneedle array 100 includes an arrayof hollowed microscale-sized needles 101, in which each needle 101comprises a protruded needle structure having an exterior wall forming ahollow interior and an opening at the terminal end of the protrudedneedle structure to expose the hollow interior, and a probe 102 formedinside the exterior wall to interact with one or more chemical orbiological substances that come in contact with the probe 102 via theopening to produce a probe signal (e.g., such as a sensing signal).

FIG. 1C shows the exemplary microneedle array 100 including an array ofwires 103 that are coupled to corresponding probes 102 of the array ofhollowed needles 101, respectively, e.g., which can provide an array ofindividually addressable microneedle sensing electrodes. The array ofwires 103 can be configured within a substrate 105, e.g., such as aninsulative material, which in some examples can be flexible and adhesiveto biological tissue. Each wire of the array of wires 103 iselectrically conductive to transmit the probe sensing signal produced bya respective probe to a sensor circuit, in which the probe sensingsignals are processed.

FIG. 1D shows an image of an exemplary microneedle imaged by scanningelectron microscopy. FIG. 1E shows an image of an exemplary microneedlearray near objects, such as a penny or an electronic circuit on aprinted circuit board, to provide size scaling of the exemplarymicroneedle array. FIG. 1F shows a zoomed view of the image of theexemplary microneedle array.

The exemplary microneedle array based sensor actuator device can involvetechniques in microfabrication, electrochemistry, enzyme-immobilizedelectrodes, and ion-selective electrodes. Potentiometric, voltammetric,amperometric, conductometric, and/or impedimetric detectionmethodologies can be integrated into one all-inclusive platform, e.g.,in order to enable the direct biosensing of multiple analytes residingin bodily fluids (e.g., such as key biomarkers occupying the transdermalfluid). For example, the microneedle array platform can be integrated onan adhesive patch that is placed on the skin in order to monitor keyphysiological and biochemical parameters transdermally. The exemplaryadhesive patch can further be integrated with electronics to allowcommunication and signal transduction. For example, because the chemicalinformation can be converted to the electrical domain viaelectrochemistry, the device can be interfaced with electronic readout,e.g., which can be analogous to continuous-monitoring blood glucosedevices. In this fashion, for example, the disclosed technology canminiaturize and integrate multiple laboratory-based tests into a singlearrayed microneedle sensing platform and provide the ability to deliveran autonomous therapeutic intervention in a controlled andminimally-invasive fashion, as well as to tailor a cocktail of drugs fordifferent forms of injury/trauma.

FIGS. 1G-1I show images of an exemplary microneedle array device on anadhesive patch implemented on living beings. FIG. 1G shows an image ofan exemplary adhesive patch employing the microneedle arraysensor-actuator device being worn on a human arm. FIG. 1H shows an imageof another exemplary adhesive patch employing the microneedle arraysensor-actuator device being worn on an animal. FIG. 1I shows anenlarged image of the exemplary device after the adhesive patch has beenremoved from the animal's skin, e.g., showing the exemplary microneedlesintact.

The disclosed biosensor-actuator technology can be used to extractrelevant physiological information and provide a controlled therapeuticresponse based on the detected physiological and biochemicalinformation. FIG. 1J shows an exemplary diagram of a “Sense-Act-Treat”feedback loop where the sensed information is used to control theactuator to adjust the drug delivery. The information from the sensingoperation enables the drug delivery to be tailored according to thesensed information. A biosensor-actuator device 180 in FIG. 1J caninclude a multiplexed array of microneedles (e.g., which can beconfigured in a manner as the exemplary microneedle array 100), in whichsome microneedles of the array are configured for sensing and othermicroneedles of the array are configured for therapeutic intervention.

The “Sense-Act-Treat” feedback loop, as shown in FIG. 1J, includes thesensing contingent of the biosensor-actuator device 180 to extract thephysiological information of an analyte from a biological fluid (e.g.,such as transdermal fluid). The exemplary sensing feature can includeindividually addressable microneedles functionalized as anelectrochemical transducer that can detect patterns of biomarkerchanges, e.g., such as acute conditions or chronic disease conditions. Aunique analyte can be detected at each microneedle within the sensingarray, or multiple analytes can be detected by several electrodes housedin one microneedle. For example, various catalysts, biocatalysts,substrates, reagents, cofactors, and/or coreagents can be immobilizedwithin the transducers to impart selectivity towards the analyte ofinterest. Likewise, ion-selective membranes (or solid state ionselective components) can be employed with electrochemical measurementsto impart selectivity towards the ions of interest. In some examples,the sensing contingent of the biosensor-actuator device 180 can alsoinclude analyte logic-gate sensing for direct processing of the sensedanalyte information, as shown in an exemplary diagram 170. For example,the diagram 170 shows two exemplary inputs (e.g., Input 1 and Input 2)of the analyte sensing logic, in which the Input 1 is a detected signalgenerated at a microneedle probe configured to detect a first reaction,and the Input 2 is a detected signal generated at another microneedleprobe configured to detect a second reaction. The diagram 170 shows thetwo exemplary input signals passed through a logic gate (e.g., a singleNAND gate in this example), in which the output of the logic gate can beused as a signal to control a microneedle actuator (e.g., in which thelogic gate output signal is interfaced with an exemplary conductingpolymer to control the porosity of a drug stored in a reservoir). Inother examples, the sensed analyte information can be processed by aprocessing unit.

For example, the input biomarkers for a soft tissue injury (STI) caninclude creatine kinase (CK) and lactate dehydrogenase (LDH), which areincident on a biocatalytic cascade, and can be representative of theInput 1 and Input 2, respectively, as shown in the diagram 170. Forexample, CK converts the creatine substrate into phosphocreatine, whichsimultaneously causes the compound to convert ATP to ADP. In thepresence of phosphoenolpyruvate (PEP), pyruvate kinase (PK) can giverise to pyruvate. If lactate dehydrogenase is present, the pyruvate canbe converted to lactate while NADH is simultaneously oxidized to NAD+.Thus, the decrease in NADH can be monitored with respect to time in anamperometric fashion. For example, since only the presence of both CKand LDH causes a concomitant decrease in NADH, monitoring Input 1 andInput 2 using the exemplary biosensor-actuator device 180 caneffectively function as a NAND Boolean logic gate.

The “Sense-Act-Treat” feedback loop, as shown in FIG. 1J, includes anexemplary image of a processing unit 175 to process the sensed analyteinformation as data and employ logic and/or instructions to control theactuator contingent of the biosensor-actuator device 180 to release, notrelease, or adjust the release of a therapeutic agent. For example, theprocessing unit 175 can include a processor and a memory coupled to theprocessor. The processing unit 175 can include a power supply, e.g.,including battery sources, renewable energy sources (e.g., solar powersources), or self-powering sources (e.g., motion feedback powersources). The processing unit 175 can include an input/output (I/O)unit, coupled to the processor and memory, which can also be connectedto an external interface, source of data storage, or display device.Various types of wired or wireless interfaces compatible with typicaldata communication standards, e.g., including, but not limited toUniversal Serial Bus (USB), IEEE 1394 (FireWire), Bluetooth, IEEE802.111, Wireless Local Area Network (WLAN), Wireless Personal AreaNetwork (WPAN), Wireless Wide Area Network (WWAN), WiMAX, IEEE 802.16(Worldwide Interoperability for Microwave Access (WiMAX)), and parallelinterfaces, can be used to implement the I/O unit. For examples, the I/Ounit of the processing unit 175 can be in communication with thebiosensor-actuator device 180 using a wired configuration. In otherexamples, the I/O unit of the processing unit 175 can include wirelesscommunication functionalities to receive data from the sensingcontingent and transmit control data to the actuator contingent of thebiosensor-actuator device 180. In such examples, the biosensor-actuatordevice 180 can include a wireless transmitter/receiver on or remotelytethered (e.g., using wires) to the substrate facilitating thesensor-actuator microneedle arrays. In such examples, the wirelesstransmitter/receiver can be interfaced with multiplexing capabilities tomultiplex the sensing signals and control signals that are transmittedand received.

The “Sense-Act-Treat” feedback loop, as shown in FIG. 1J, includes theactuator contingent of the biosensor-actuator device 180 to deliver oneor more drugs to the region penetrated by the microneedles based on theprocessed analyte information (sensed by the sensing contingent). Theexemplary drug delivery feature can enable the autonomous delivery of atargeted therapeutic intervention in response to the detected acute orchronic condition. For example, the permeability of the conductingpolymer nanoactuators can be tunable through an autonomous porositychange controlled by the integrated sensing or enzyme logic system(e.g., processed by the processing unit 175), which in turn can controlrelease of the drug, as illustrated in a diagram 185 of the drugdelivery actuator in FIG. 1J. The diagram 185 shows that a processedsignal represented by 0 does not actuate the release of the drug (e.g.,the porosity of the exemplary conducting polymer remains in aneffectively closed stat). The diagram 185 also shows that a processedsignal represented by 1 does actuate the release of the drug (e.g., theporosity of the exemplary conducting polymer is triggered to be in anopen state, thereby allowing the drug to pass through the pores of thepolymer and exit the microneedles into the tissue fluid). The arrayedmicroneedle structure can allow multivariate/multiplexed drug delivery,and a unique therapy can be delivered at each microneedle constituent ofthe array, as shown in FIGS. 1K and 1L.

FIG. 1K shows a diagram of the exemplary microneedle arraysensor-actuator 180 featuring individually-addressable microneedles tomonitor for different dysfunctions and individually-addressablemicroneedles to deliver a therapeutic agent in response to detection ofthe dysfunctions. The diagram of the biosensor-actuator device 180, asshown in FIG. 1K, includes microneedles of the array that are configuredfor sensing (e.g., microneedles 181 a, 181 b, and 181 c) and othermicroneedles of the array that are configured for therapeuticintervention (e.g., microneedles 181 d, 181 e, and 181 f). In thisexample, the microneedle 181 a is configured for sensing an analyteassociated with soft tissue injury (STI), the microneedle 181 b isconfigured for sensing an analyte associated with traumatic brain injury(TBI), and the microneedle 181 c is configured for sensing an analyteassociated with abdominal trauma (ABT). Also in this example, themicroneedle 181 d is configured for delivering a drug associated withtreating STI, the microneedle 181 e is configured for delivering a drugassociated with treating TBI, and the microneedle 181 f is configuredfor delivering a drug associated with treating ABT.

For example, exemplary analytes associated with STI that can be detected(e.g., using the sensor-actuator 180) include creatine kinase, lactate,and lactate dehydrogenase; e.g., ameliorated with glucocorticoids,NSAIDs. Exemplary analytes associated with TBI that can be detectedinclude glutamate, ceruloplasmin; e.g., ameliorated with acetaminophen.Exemplary analytes associated with ABT that can be detected includelactate, lactate dehydrogenase; e.g., ameliorated with acetylsalicylicacid or iso-butyl-propanoic-phenolic acid.

FIG. 1L shows an illustrative diagram of a multiplexed controlledrelease of a targeted therapeutic cocktail, e.g., where the polymeractuator(s) is individually addressable for on-demand release fortargeted therapeutic intervention. The diagram of FIG. 1L showsexemplary microneedles of the actuator contingent (e.g., microneedles182 d, 182 e, and 182 f) that are configured for therapeuticintervention. In this example, the microneedle 182 d is configured forcontrolled release of a drug 183 d that is stored in Reservoir 1, themicroneedle 182 e is configured for controlled release of a drug 183 ethat is stored in Reservoir 2, and the microneedle 182 f is configuredfor controlled release of a drug 183 f that is stored in Reservoir 3.For example, based on a control signal received from the processing unit(e.g., like that shown in the diagram 185 in FIG. 1L), the release ofany or all of drugs 183 d, 183 e, and/or 183 f can be controlled (e.g.,using multiplexing) to produce a targeted therapeutic cocktail, e.g., inwhich the individually-addressable polymer actuator(s) are actuated in amanner that can control the size of the porosity (e.g., and thereby theflow), as well as the duration of the open state, controllingconcentration of each of the released drugs.

FIGS. 2A-2C show schematics of an exemplary microneedle array stripsystem 200. In this example, the system can include a flat flex cable(FFC) 210 that includes a plurality of conductors 211 (e.g., ten copperconductors) that interconnect the microneedle array to a connectorregion 218, e.g., in which the conductors 211 can be 1.5″ length andinterface (e.g., right end (re)connected) to a circuit board, e.g., viaa zero insertion force (ZIF) connector. For example, the exemplarymicroneedle array strip system can be configured on a FFC sized to be11.0 mm by 38.1 mm. Holes can be opened using laser ablation at the leftend to expose underlying traces. Metal electrodes, e.g., four workingelectrodes 202 and one counter electrode 207 and one or more referenceelectrode(s) 206, can be sputter deposited on the surface over theopenings. Four vented fluidic chambers 215 can be made from laserablated Mylar with a pressure sensitive adhesive. The adhesive layer canalso bond microneedle array(s) to the FFC. FIG. 2A shows a top view ofthe schematic of the microneedle array strip system 200. FIG. 2B shows athree dimensional view of the schematic of the microneedle array stripsystem 200. FIG. 2C shows another three dimensional view of theschematic of the microneedle array strip system 200 inserted into acircuit board connector 219.

In another example, individually addressable electrodes (microneedles)can be loaded with a carbon paste, carbon fiber, or conducting polymertransducer and be employed for the detection of patterns of biomarkerchanges that reflect optimal health and/or performance. Usingpotentiometry, amperometry, or voltammetry, various catalysts,biocatalysts, substrates, reagents, cofactors, and/or coreagents can beimmobilized within the transducers to impart selectivity towards theanalyte of interest. Likewise, ion-selective membranes (or solid stateion selective components) can be employed with electrochemicalmeasurements to impart selectivity towards the ions of interest.Significant predictive and diagnostic information can be available inmonitoring multiple biomarkers and in measuring the dynamical pattern ofthose species as a measure of the overall health/performance/fitness ofthe subject. Patterns in multiple biomarkers can be integrated andchanges in those markers can be assessed over extended time periods inorder to provide a more detailed and accurate temporal characterizationof the negative effects of stress and overtraining in addition to aplethora of diseases and illnesses.

For example, arraying the microneedles can allow for measuring patternsin multiple bioanalytes. Moreover, an analyte or multiple analytes, suchas a catalyst/biocatalyst or other analyte or biomarker substance, canbe immobilized by robust means that can includeelectropolymerization/polymer entrapment, electrostatic interactions,covalent attachment, and direct adsorption. In one example, a planarsolid-state transducer can be an electrode, for example, use of carbonfiber, carbon paste, and conducting polymers to form the electrochemicaltransducer.

The exemplary device can be utilized in the following manner. Atransdermal microneedle array can be employed; each microneedleconstituent can contain a bored cylindrical vacancy inside which athree-electrode electrochemical sensing element is housed (e.g., such aspotentiometric, voltammetric, amperometric, conductometric,impedometric, etc. sensing elements). In one example, an enzyme (withaffinity to a particular biochemical moiety) can be immobilized on theworking electrode of the three-electrode contingent and amperometry canbe performed. In another example, an ion-selective membrane (withsuitable ionophore) or solid state functionalization can be applied tothe working electrode and potentiometry or voltammetry is performed. Thepresence of the analyte, metabolite, electrolyte, or ion of interest canresult in perturbations in the detected current (enzyme electrode) orpotential (ion-selective electrode), respectively.

FIGS. 3A-3C show data plots of exemplary results of multiplexeddetection of various biomarkers on an electrochemical array strip usingindividual microneedles. FIG. 3A shows a data plot of multiplexeddetection of pH; FIG. 3B shows a data plot of multiplexed detection oflactate; and FIG. 3C shows a data plot of multiplexed detection ofglucose.

FIGS. 4A and 4B show exemplary individual microneedle microsensors thatcan be addressable through a microelectrode array mated with the reverseside of the device. For example, the array(s) can be fabricated using anumber of techniques including photolithography, inkjet printing, andscreen printing, among other techniques. An example of a screen printedmicroneedle array is shown in FIG. 4A, and a microneedle arrayfabricated by a stenciling process used to define the pattern is shownin FIG. 4B.

In another embodiment of the disclosed technology, a minimally-invasivemulti-component microneedle device for electrochemical monitoring andbiosensing is described. This embodiment can comprise the sameembodiment(s) like those previously described, and can thereforeimplement the entirety of functionalities of the individual embodimentson a single embodiment. For example, the disclosed technology can beimplemented for the electrochemical monitoring and biosensing of theexcitatory neurotransmitter glutamate and glucose. In this exemplaryembodiment, a device can include tight integration of solid and hollowmicroneedles into a single biosensor array device containing multiplemicrocavities. Such microcavities can facilitate the electropolymericentrapment of the recognition enzyme within each microrecess. Theresulting microneedle biosensor array can be employed as an on-bodyminimally-invasive transdermal patch, e.g., eliminatingextraction/sampling of the biological fluid, thereby simplifying devicerequirements.

Exemplary implementations were performed to demonstrate variousfunctionalities of the device, e.g., including the electropolymericentrapment of glutamate oxidase and glucose oxidase within apoly(o-phenylenediamine) (PPD) thin film. For example, the PPD-basedenzyme entrapment methodology can enable the effective rejection ofcoexisting electroactive interferents without compromising thesensitivity or response time of the device. The resultingmicroneedle-based glutamate and glucose biosensor can exhibit highselectivity, sensitivity, speed, and stability in both buffer anduntreated human serum. For example, high-fidelity glutamate measurements(e.g., down to the 10 μM level) were obtained in undiluted human serum.The exemplary recess design can also protect the enzyme layer uponinsertion into the skin. The described robust microneedle design can bewell-suited for diverse biosensing applications in which real-timemetabolite monitoring is a core requirement.

The exemplary microneedle-based glutamate and glucose biosensor wereimplemented in ways that demonstrate clinical application ofmicrodevices for the on-body monitoring of relevant bioanalytes byminimally-invasive electrochemical biosensors. In this regard, themicroneedle arrays can be configured to provide pain-free biosensingand, being highly integrated biocompatible devices, these devices can befabricated on an industrial scale and at low cost. For example, thedescribed microneedle arrays can perform monitoring and biosensingapplications without involving fluid sampling/extraction. For example, afeature of the technology is that the uptake of bodily fluids (e.g.,such as transdermal fluid) is not required, as is in conventionalmicrofluidic sensing systems. Through the execution of electrochemistryat the microneedle-transdermal fluid interface, useful chemicalinformation can be extracted and directly transduced to the electronicdomain. In this manner, sophisticated and costly mechanical devices thatregulate flow to a detector array/separate sensing unit can beeliminated from implementation.

The exemplary microneedle sensing array device can employ a recess-basedmicrocavity structure that can be designed to confine the recognitionenzyme and protect it upon penetration of the skin. For example, abicomponent microneedle biosensor can include an array ofplatinum-coated solid microneedles, which can serve as the workingelectrode, and a hollow microneedle cover, which can provide amicrocavity that surrounds each solid microneedle. This exemplarybicomponent microneedle biosensor is shown in FIGS. 5A-5D. FIGS. 5A-5Dalso illustrate a process to fabricate the exemplary bicomponentmicroneedle biosensor.

FIG. 5A shows a solid microneedle constituent 501 and a hollowmicroneedle constituent 503 of an exemplary array of bicomponentmicroneedle electrodes 500. The solid microneedle constituent can becoated with a conductive material to form a working electrode, e.g.,such as a platinum working electrode 502. FIG. 5A also shows a processto assemble the platinum-coated solid microneedles 502 with the hollowmicroneedle cover 503. In some examples, the assembly includes applyinga sealing agent (e.g., epoxy) to the non-detecting, non-recessed regionsbetween the solid microneedle constituent 501 and the hollow microneedleconstituent 503. FIG. 5B shows the bicomponent microneedle arrayelectrode 500 fully-assembled. As shown in FIG. 5B, each bicomponentmicroneedle electrode 500 in the array includes a recess region thatexposes the needles (e.g., the platinum-coated electrodes 502) of thesolid microneedle constituent 501, e.g., such that the recess canfacilitate enzyme immobilization. FIG. 5C shows the growth of theglutamate oxidase (GluOx)-functionalized PPD film 504 at the solidmicroneedle surface within the recess region of the microneedleelectrode 500 from the o-phenylenediamine (o-PD) monomer of a o-PD-GluOxsolution. For example, fabrication of the array of bicomponentmicroneedle electrodes 500 can include applying a GluOx-PPD thin film tothe platinum working electrodes 502 by immersing the microneedles in ano-PD-GluOx solution. FIG. 5D shows the biocatalytic behavior of theelectropolymerized glutamate oxidase-poly(o-phenylenediamine) film 504(illustrated in purple), e.g., enabling the quantification of glutamatelevels within the transdermal fluid. In these exemplary figures, glucoseoxidase (GOx) becomes substituted in place of GluOx for thequantification of glucose.

The recess-based microneedle electrodes 500 can enable electropolymericentrapment of the enzyme within the individual microcavities. As aresult, direct transdermal biosensing can be accomplished withoutrequiring the uptake of the transdermal fluid, e.g., thereby simplifyingdevice requirements and the sensing process. The bicomponent recessgeometry of the microneedle biosensor can also provide a greater surfacearea for enzyme immobilization with microneedles containing embeddedplanar electrodes.

The recess-based microneedle electrodes 500 can includeelectropolymerized PPD thin films that can be employed for theconfinement of enzymes into miniaturized electrode transducers, e.g.,while imparting remarkable permselective properties and a stableresponse. In one example, PPD can be used for entrapping differentoxidase enzymes such as glucose oxidase, lactate oxidase, and glutamateoxidase, along with permselective detection of the liberated hydrogenperoxide product. As a consequence of their remarkable permselectiveproperties, PPD films can impart high selectivity and stability throughexclusion of co-existing electroactive interferences and proteinsnormally present within bodily fluids. The described biosensor devices,which can employ PPD films, can thus facilitate the amperometricdetection of hydrogen peroxide with high substrate selectivity,excellent sensitivity, operational stability, and rapid response time.In this manner, the described biosensor devices that employenzyme-functionalized PPD films can exhibit considerable sensingadvantages when compared with those based on other immobilizationtechniques, as described herein.

For example, to illustrate the versatility of the disclosed bicomponentmicroneedle array platform, an exemplary biosensor device foramperometric glutamate biosensing is demonstrated. This exemplaryplatform can be subsequently extended to glucose monitoring for themanagement of diabetes mellitus. For example, an excitatoryneurotransmitter, glutamate, can be implicated in a number of pathologicmedical conditions such as ischemic neuronal injury, hypoglycemicinjury, epilepsy, Alzheimer's disease, and traumatic brain injury. Inaddition, elevated glutamate levels in the circulatory system can beassociated with excitotoxicity. Blood glutamate levels have risen froman average value of 37.5 μM among healthy patients to 141.3 μM amongpatients who have sustained moderate to severe trauma related tointracranial injury. As such, serum glutamate levels can provide usefulinsight into the overall condition of the central nervous systemfollowing brain trauma.

For example, the described biosensing platform can be advantageous overbiosensors that quantify glutamate levels with a high degree ofinvasiveness, e.g., such as by uptaking the cerebrospinal fluid (CSF)via a catheter or a microdialysis probe for further analysis. Also forexample, the described biosensing platform can be advantageous overbiosensors that typically are clinically implemented in a hospitalsetting, e.g., as such clinical analysis can be a painful,time-consuming, and costly proposition. In addition, the describedbiosensing platform can be amenable to on-body continuous monitoring,especially when access to the CSF is not feasible. As blood glutamatelevels correlate well with the levels found in the CSF, its extractionfrom this hard-to-access bodily fluid is unnecessary under the disclosedembodiment.

An exemplary demonstration of the described biosensor device involvesthe enzymes glutamate oxidase (GluOx) and glucose oxidase (GOx) that canbe entrapped within the microcavities of the exemplary microneedledevice using different PPD growth processes, each with its own specificadvantage that can be tailored to specific applications. The PPD-basedconfinement of the enzymes within the microneedle cavities can enablethe efficient quantification of glutamate and glucose atpathophysiological levels within buffer solutions and undiluted humanserum. For example, the minimally-invasive nature of the exemplarydevice, combined with its convenient means to achieve enzyme entrapmentand protection, as well as its attractive electroanalytical performance,can demonstrate its applicability as a practical patch-type on-bodybiosensor.

Exemplary materials and methods to implement the disclosed embodiment ofthe technology are presented. The following chemicals and reagents wereused in the described implementations, which included glutamate oxidase(GluOx, E.C. 1.4.3.11) from E. coli (recombinant), glucose oxidase (GOx,E.C. 1.1.3.4) from Aspergillus niger, 1,2-phenylenediamine (o-Pd),L-glutamatic acid (GLU), D-(+)-glucose (GLC), L-ascorbic acid (AA), uricacid (UA), L-cysteine (CYS), acetaminophen (ACT), sodium sulfate,ethylenediaminetetraacetic acid (EDTA), potassium phosphate monobasic,potassium phosphate dibasic, and serum from human male (type AB). Theexemplary implementations (with the exception of the serum calibration)were performed in 0.1 M phosphate buffer (pH 7.40) with 0.5 mM EDTA.Ultrapure water (18.2 MΩ·cm) was employed in all exemplaryimplementations.

The instrumentation used in the described implementations included thefollowing, which was utilized in exemplary demonstrations andimplementations of the disclosed embodiment under exemplary conditionsdisclosed herein. A CH Instruments (Austin, Tex.) model 1232Aelectrochemical analyzer was employed for electrochemical measurements.An external Ag/AgCl reference electrode (CH Instruments CHI111) and a0.5 mm diameter platinum wire counter electrode (BASi, West Lafayette,Ind.) were used to establish a three-electrode electrochemical system.Voltammetric and chronoamperometric studies were used to evaluate theelectrochemical behavior of the microneedle array electrode at roomtemperature (22° C.). In these exemplary electrochemicalimplementations, glutamate (or glucose) was added into 2 mL of phosphatebuffer solution or serum (stirred) in order to obtain the desiredconcentration. Chronoamperometric currents were sampled for 15 sfollowing the application of the potential step. The morphology of thebicomponent microneedle array was examined using a field emissionscanning electron microscope (SEM) (Philips XL30, Amsterdam, theNetherlands). Specimens were coated with chromium prior to SEM analysisusing a sputtering instrument (Energy Beam Sciences Emitech K575X, E.Granby, Conn.). A deposition current of 130 mA was applied for 30 s todeposit ˜15 nm of chromium on the sample surface.

The exemplary solid and hollow microneedle arrays used in the exemplaryimplementations were developed in the following manner. The microneedledesigns were prepared using a CAD software, e.g., Solidworks (DassualtSystemes S.A., Velizy, France). Substrate support structures werecreated with Magics RP 13 (Materialise NV, Leuven, Belgium). Forexample, the solid needles were designed and fabricated with a conicalin shape and possess a base diameter of 390±14 μm and a height of 818±35μm. The hollow needles were pyramidal in shape with a triangular base.The dimensions of each hollow microneedle were as follows: an edgelength of 1174±13 μm, a height of 1366±15 μm, and a vertical cylindricalbore of 342±5 μm diameter on one of the faces of the pyramid structure.Both the solid and hollow needles were arranged into 3×3 square arrayswith 2 mm periodicity. Substrates for the microneedle arrays were 10mm×10 mm in extent and possessed thickness values of 2000 μm and 500 μmfor solid and hollow variants, respectively. The three-dimensionalcomputer models were transferred to a Perfactory® SXGA Standard UV rapidprototyping system (EnvisionTEC GmbH, Gladbeck, Germany) forfabrication. This system used computer models to precisely guide lightfrom a 150 W halogen bulb over a photocurable material, e.g., resultingin the selective polymerization of the exposed material. In someaspects, Eshell 200 acrylate-based polymer (EnvisionTEC GmbH) can beutilized as the constituent material to fabricate the microneedle arrayssince the resin selectively polymerizes under visible light and exhibitsa Young's modulus of elasticity of 3050±90 MPa. Moreover, the polymerfeatures Class-Ha biocompatibility per ISO 10993. A 550 mW output powerbeam (step size=50 μm) with a zero-degree tilt was employed for thepolymerization of the resin. Following the fabrication routine, thearrays were rinsed with isopropanol to remove the unpolymerizedmaterial. The arrays were placed in an Otoflash post curing system(EnvisionTEC); and post-build curing was performed for 50 s. A Compex201 krypton-fluoride (KrF) excimer laser (Coherent, Santa Clara,Calif.), which can be operated with a 10 Hz repetition rate and awavelength of 248 nm, was used to ablate a commercially-obtained highpurity Pt target. This process resulted in the deposition of thin filmsof Pt (˜12 nm) on the surface of the solid microneedle array. Abackground pressure of 5 μTorr was maintained during the 2 min pulsedlaser deposition (PLD) routine, performed at room temperature.

Adhesive non-conducting epoxy can be applied to the periphery of thesolid microneedle substrate. The hollow microneedle cover can then beplaced over the solid microneedle substrate. This exemplary procedure isdiagrammatically represented in FIGS. 5A and 5B. For example, the twocomponents (e.g., the solid microneedle substrate and the hollowmicroneedle cover) can be arranged under an optical microscope to alignthe solid microneedles within the hollow microneedle aperture. This canform the bicomponent microneedle array electrode (BMAE), as shown inFIG. 5C, e.g., which can be mated with a 3 mL syringe (BD Biosciences,Franklin Lakes, N.J.). For example, the nozzle portion of the syringecan be removed to facilitate the attachment of the BMAE, which can beaffixed using adhesive epoxy to the syringe tip for easier handling. Acopper wire can be subsequently inserted into the open end of thesyringe in order to create an electrical contact to the Pt workingelectrode. A poly(o-phenylenediamine) (PPD) film can beelectropolymerized from a solution of the o-phenylenediamine (o-Pd)monomer, as shown in FIG. 5C, e.g., to immobilize the GluOx and GOxenzymes on the electrode surface and reject potential electroactiveinterferents. For example, a 0.1 M phosphate buffer (pH 7.40) solutioncontaining 10 mM o-Pd, 5 mM sodium sulfate, and 100 U/mL GOx can bepurged with nitrogen for 20 minutes at room temperature, which can beused to form the GOx-functionalized electrode. For example, the BMAE,Ag/AgCl reference, and platinum counter electrodes can then be immersedin the solution; a potential of 0.75 V vs. Ag/AgCl can subsequently beapplied for 20 min in order to grow the GOx-entrapped PPD film, asrepresented in FIG. 5C. This exemplary process represents a rapid meansto immobilize enzymes and is appropriate for applications in which theenzyme is of sufficiently low cost such that the entire extent of theelectrode can be immersed in the enzyme-o-PD solution.

In other examples, a slight variant of the aforementioned process can beused, e.g., to conserve the costly GluOx enzyme during theelectropolymerization process. In this alternative exemplary process,the BMAE, Ag/AgCl reference, and platinum counter electrodes can beimmersed in a solution of 0.1 M phosphate buffer (pH 7.40) containing 10mM o-Pd and 5 mM sodium sulfate; and a potential of 0.75 V vs. Ag/AgClcan subsequently be applied for 5 min. The electrode can then be rinsedand dried at room temperature. A 0.5 μL aliquot solution of 7.5 U/mLGluOx in 0.1 M phosphate buffer (pH 7.40) can then be dispensed in eachrecess of the BMAE; this step can be repeated an additional six times oneach microneedle using a low-retention micropipette. Following thisprocess, the solution can be allowed to dry at room temperature. Thedrop-casting procedure can be repeated five additional times.Subsequently, a solution of 0.1 M phosphate buffer (pH 7.4) containing10 mM o-Pd, 5 mM sodium sulfate, and 1 U/mL GluOx can be dispensed oneach microcavity of the microneedle array. A potential of 0.75 V vs.Ag/AgCl can subsequently be applied for 15 min in order toelectropolymerize the GluOx-entrapped-PPD film. For example, whereas theprevious described process can be adapted for simplicity at the expenseof increased enzyme usage, this implementation can facilitate theelectropolymerization of more costly enzymes.

Following each electropolymerization process, the BMAE can be rinsed andimmersed in a 0.1 M phosphate buffer solution (pH 7.4) for 30 min toremove monomeric residue from the microneedle structure as well as anynon-bound enzyme. When not in use, the BMAE can be stored in phosphatebuffer at 4° C. This exemplary process, which is diagrammaticallyrepresented in FIG. 5D, can enable the quantification of glutamate andglucose, as well as other analytes.

The exemplary implementations of the described microneedle arraybiosensor device included an evaluation of the surface morphology of theBMAE which was performed to ascertain the electrode geometry and surfacefeatures. A close examination of the BMAE surface morphology wasexecuted using SEM. FIG. 6A shows exemplary scanning electronmicrographs of the solid microneedle arrays. FIG. 6B shows exemplaryscanning electron micrographs of the hollow microneedle arrays. As shownin the SEM images, the features of the exemplary microneedles closelycorresponded with those specified in the exemplary computer-aided designfile. For example, a uniform distribution of the solid and hollowmicroneedles can be observed at the microneedle array. With respect tothe exemplary solid microneedle (FIG. 6A), a notable observation is theribbed structure, which can be attributed to the layer-by-layer approachthat was utilized to polymerize the E-shell 200 resin. A uniformpyramidal structure and a triangular base can be observed at eachcomponent of the hollow microneedle array (FIG. 6B), along withapertures of uniform size distribution. Minimal to nomicroneedle-to-microneedle geometric variation was observed.

The exemplary implementations of the described microneedle arraybiosensor device included electrochemical characterizations of thebicomponent microneedle electrode towards the amperometric detection ofglutamate. The initial electrochemical characterization of the BMAE wasaimed at constructing hydrodynamic voltammograms (HDVs), e.g., to selectan optimal detection potential. For example, an HDV can be obtainedusing chronoamperometry at varying potentials between 0.1 and 0.6 V vs.Ag/AgCl (in 50 mV increments). These exemplary characterizations wereperformed in the blank buffer solution containing 100 μM of glutamate.The redox currents was sampled at 15 s following the potential step. Anidentical procedure can be followed using GOx-BMAE for the detection of10 mM glucose. FIG. 7 shows exemplary hydrodynamic voltammograms of theglutamate bicomponent microneedle array electrodes, e.g., in which redoxcurrents were sampled at t=15 s following the potential step. As isevident from the figure, the presence of glutamate caused a concomitantrise in the anodic current, corresponding to the oxidation of the H₂O₂enzymatic product. The onset of the peroxide oxidation can occur at˜0.25 V vs. Ag/AgCl. To minimize the potential oxidation of interferentsin real samples, a potential of 0.40 V vs. Ag/AgCl can be selected forfurther electrochemical implementations of the BMAE biosensor.

Examples are described for biosensing of glutamate in a buffer matrixand human serum at the bicomponent microneedle array electrode. Thesensitivity of the exemplary BMAE biosensor was evaluated usingchronoamperometric potential steps to the selected potential of 0.40 Vvs. Ag/AgCl. FIG. 8A shows a data plot of the average of triplicatechronoamperometric experiments for increasing levels of glutamate overthe entire pathophysiological range (e.g., 0-140 μM in 20 μM increments,sampled at t=15 s) in a buffer matrix. A linear calibration plot(R²=0.995) can be observed (as shown in the inset of FIG. 8A) over theentire range under investigation. The exemplary calibration plotexhibits high sensitivity (s_(x)=7.129 nA/μM) and a low deviation(RSD=3.51%); a limit-of-detection (LOD) of 3 μM can be estimated basedon the signal-to-noise characteristics of the experimental data (S/N=3).The LOD lies well below normal physiological levels, reflecting theability of the microneedle sensor to detect physiologic levels ofglutamate.

Following the calibration experiments in the buffer solution, exemplaryimplementations of the described BMAE biosensor was evaluated byquantification of glutamate in undiluted human serum samples. FIG. 8Bshows a data plot of the average of triplicate chronoamperometriccalibration experiments in human serum for increasing levels ofglutamate over the 0-140 μM range (e.g., in 20 μM increments). As withthe exemplary buffer study, a linear calibration plot can be observed(as shown in the inset of FIG. 8B; R²=0.992) over the entire range. Inaddition, for example, the calibration data exhibit high sensitivity(s_(x)=8.077 nA/μM) and low deviation (RSD=6.53%). An LOD of 10 μM canbe estimated based on the signal-to-noise characteristics of these data(S/N=3). As with the exemplary buffer experiments, the LOD obtained fromthe experimental data resides below the limit of normal physiologicallevels. The attractive behavior of the BMAE in untreated serum samplesreflects the protective ability of the PPD film. Furthermore, thesimilar sensitivity obtained for both the buffer- and serum-based trialsagain can underscore the robustness of the PPD immobilization schemedespite the prolonged exposure of the biosensor to protein-rich serummedium.

Examples are described for interference investigation(s) employingphysiologically-relevant electroactive compounds. For example, anotheradvantageous feature of the PPD coating includes its ability to rejectcoexisting electroactive interferents even at moderate oxidationpotentials. Accordingly, the selectivity of the response was examined inthe presence of physiological levels of ascorbic acid (60 μM), uric acid(500 μM), cysteine (200 μM), and acetaminophen (200 μM). FIG. 9 shows adata plot of exemplary chronoamperograms recorded in 0.1 M phosphatebuffer (pH 7.40) for the blank buffer solution, 100 μM glutamate, and100 μM glutamate in the presence of the electroactive physiologicalinterferents ascorbic acid (AA, 60 μM), uric acid (UA, 500 μM), cysteine(CYS, 200 μM), and acetaminophen (ACT, 200 μM). The exemplaryimplementations were carried out under the same conditions as thoserepresented in FIGS. 8A and 8B. FIG. 9 illustrates the negligiblecontribution imparted by the presence of these electroactive compoundsupon the current signal for 100 μM glutamate. Physiological levels ofascorbic acid, uric acid, cysteine, and acetaminophen resulted in only0.44%, 0.31%, 1.93%, and 6.37% average deviations from the 100 μMglutamate current response, respectively. Hence, natural metabolicfluctuations in the levels of these electroactive species are notexpected to interfere with the in vivo quantification of glutamate usingthe exemplary BMAE device as an on-body biosensor.

Exemplary implementations were performed for stability analysis of theexemplary bicomponent microneedle electrode array. For example, theability to operate over prolonged periods with minimal deterioration inthe current response can represent another important feature of on-bodybiosensors. Accordingly, the stability of the BMAE response was examinedusing a 100 μM glutamate solution over an eight-hour period. FIG. 10shows a data plot showing stability of the glutamate response overextended time periods with each data item referenced to the originalcurrent level at t=0 (100%). In this exemplary implementation, data wasgenerated from chronoamperograms recorded in 0.1 M phosphate buffer (pH7.40 ) with 140 μM glutamate. The exemplary implementations were carriedout under the same conditions as those represented in FIGS. 8A and 8B.The exemplary time-course profile of FIG. 10 indicates that theexemplary biosensor exhibits a highly stable current response, retaining105% of the original signal level after eight hours of continuoussampling. In this example, the measured current never exceeded 110% ofthe original level over the entire time period. Consequently, the BMAEcan be anticipated to perform reliably over extended periods associatedwith body-worn biosensors.

Examples are described for the biosensing of glucose with the exemplarybicomponent microneedle electrode array. For example, with the glutamateBMAE methodically characterized, the exemplary platform can besubsequently migrated for use as a glucose biosensor towards diabeticmonitoring. GOx was confined within the BMAE cavities using a slightvariant of the described GluOx immobilization process. This techniquecan be amenable to cost-per-quantity considerations, as stated above.FIG. 11A shows a data plot featuring chronoamperometric calibration datafor increasing levels of glucose over the entire pathophysiologicalrange (e.g., 0-14 mM in 1 mM increments) in a buffer matrix. Awell-defined response can be observed over the entire range, leading toa linear calibration plot (as shown in the inset of FIG. 11A; R²=0.996).In addition, these exemplary calibration data exhibit high sensitivity(s_(x)=0.353 μA/mM) and low deviation (RSD=6.44%, n=3), along with a LODof 0.2 mM (S/N=3). It can be noted that the GOx-functionalized BMAEexhibited a lower sensitivity towards its substrate when compared withthe GluOx-functionalized platform. For example, this can be attributedto the different PPD growth process, which may affect the transportproperties of the substrate and product. Accordingly, the GluOximmobilization process can be followed when high sensitivity is desired.

FIG. 11B demonstrates the high selectivity of the exemplary glucosemicroneedle biosensor. FIG. 11B shows a data plot of exemplarychronoamperograms recorded for the blank buffer solution, 10 mM glucose,and 10 mM glucose in the presence of the electroactive physiologicalinterferents ascorbic acid (AA, 60 μM), uric acid (UA, 500 μM), cysteine(CYS, 200 μM), and acetaminophen (ACT, 200 μM). FIG. 11B illustrates thecontribution imparted by the presence of potential electroactiveinterferents upon the current signal for 10 mM glucose. Physiologicallevels of ascorbic acid, uric acid, cysteine, and acetaminophen resultedin negligible deviations of 1.07%, 0.88%, 1.65%, and 2.21%,respectively, from the current response for 10 mM glucose. Consequently,as with the exemplary interference implementation conducted with theglutamate BMAE, natural metabolic fluctuations of these compounds is notbe anticipated to interfere with the monitoring of glucose.

An exemplary stability evaluation of the GOx-functionalized BMAE wasperformed using a buffer solution containing 10 mM glucose over an8-hour period. The GOx BMAE yields a highly stable current response,with 97% of the original signal level extracted at the conclusion of themeasurement period. In this example, throughout the time period underinvestigation, the measured current response never fell below 87% of theoriginal level. The similar results described earlier with theGluOx-functionalized BMAE indicate that both PPD-based immobilizationschemes can yield a stable response over a prolonged period ofcontinuous use.

The disclosed technology described for this embodiment includes abicomponent microneedle array biosensor platform that can be used forminimally-invasive glutamate and glucose quantification. The bicomponentmicroneedle design merges the inherent advantages of solid and hollowmicroneedles in order to form a microcavity that can allow theelectropolymeric entrapment of an enzyme, which can provide protectionfor the enzyme layer upon skin penetration, and can eliminate the needfor the extraction of biological fluids. Using apoly(o-phenylenediamine) thin-film for entrapping the enzymes glutamateand glucose oxidase can enable the highly sensitive, selective, stable,and rapid electrochemical detection of glutamate and glucose,respectively. The high fidelity detection of glutamate in undilutedhuman serum samples over the entire pathophysiological range can furthersubstantiate the utility of the platform as a practical on-bodybiosensor. The exemplary patch-type on-body biosensor can enables thetransdermal monitoring of a number of relevant metabolites.

In another embodiment of the disclosed technology, a minimally-invasive,multiplexed, multi-component microneedle actuator device that enablesthe controlled delivery of multiple therapeutic agents is described.This embodiment can comprise the same embodiment(s) like thosepreviously described, and can therefore implement the entirety offunctionalities of the individual embodiments on a single embodiment. Inthis described embodiment, a device can deliver a drug in response toinjury/trauma in an autonomous, minimally-invasive, and controlledmanner that leverages microneedle arrays as the delivery structure,which can be referred to as a smart NanoPharmacy-on-A-Chip. For example,microneedle array(s) can be integrated on an adhesive patch that isplaced on the skin in order to deliver on demand a targeted therapeuticintervention transdermally. The exemplary technology can integrate themicroneedle platform with stimuli-responsive conducting polymernanoactuators (with tunable permeability through an autonomous porositychange controlled by an integrated sensing or enzyme-logic system). Forexample, the disclosed biosensor-actuator device can be used to aid inthe rapid administration of multiple therapeutic agents and counteractdiverse biomedical conditions.

The described embodiment includes multiple individually-addressablechannels on a single microneedle array, each paired with its ownreservoir and conducting polymer nanoactuator, which are used to delivervarious permutations of the multiple unique chemical species. Forexample, upon application of suitable redox potentials to the selectedactuator, the conducting polymer is able to undergo reversible volumechanges, thereby serving to release a model chemical agent in acontrolled fashion through the corresponding microneedle channels.Exemplary implementations of the drug delivery contingent of thedisclosed biosensor-actuator device were performed and are describedherein. For example, time-lapse videos were recorded and can offerdirect visualization and characterization of the membrane switchingcapability, and, along with calibration investigations, confirmed theability of the device to alternate the delivery of multiple reagentsfrom individual microneedles of the array with high precision andtemporal resolution. Analytical modeling is described herein, which canoffers prediction of the volumetric flow rate through a singlemicroneedle, and accordingly, can be used to assist in the design of themicroneedle arrays.

In some examples, conducting polymers such as polypyrrole (PPy),polyaniline (PANI), and poly(3,4-ethylenedioxythiophene) (PEDOT) can beemployed for utilization in the described controlled release systems anddrug delivery actuators. These exemplary materials include uniqueproperties (e.g., PPy in particular), which include their reversiblemechanical behavior as “artificial muscles” and their ability to changeporosity and undergo volume changes in response to appliedelectrochemical stimuli. The disclosed actuator technology engineersthese exemplary materials in devices and systems to provide a means todeliver medications in an effective and minimally-invasive manner, e.g.,which can be implemented in practical body-worn devices for theamelioration of disease and injury in the acute phase, amenable toextended durations of pain-free wear.

FIG. 12A shows a schematic illustration of an exemplarymicroneedle-based multi-channel, multiplexed drug delivery actuatordevice 1200. The device 1200 includes a hollow microneedle array 1201.The device 1200 includes a gold-sputtered polycarbonate membrane 1204,which can be functionalized with sodium dodecylbenzenesulfonate-dopedpolypyrrole (PC/Au/PPy/DBS). The device 1200 includes apolydimethylsiloxane (PDMS) reservoir 1207 that can include multiplereservoirs to store chemical (therapeutic) agents.

FIG. 12B shows a schematic illustration of the assembledmultiple-channel drug delivery actuator device 1200, in which thereservoir 1207 is configured to have two exemplary reservoirs containingtwo different drugs, e.g., drug 1211 and drug 1212. The schematic inFIG. 12B also shows the assembled multiple-channel drug deliveryactuator device 1200 including electrode connections 1216 and 1217corresponding to the two reservoirs containing the drug 1211 and 1212,respectively.

The disclosed biosensor-actuator device enables the controlled andswitchable delivery of multiple therapeutic agents. Exemplaryimplementations of the described device utilized still-frame imaging andreal-time video capture to show the alternating release of dye fromdifferent microneedles from the same array platform. For example, imageanalysis software (e.g., such as ImageJ) and ultraviolet-visible(UV-Vis) spectrophotometry were employed to demonstrate the switchingaccuracy and repeatability of the microneedle volumetric flow rate.These exemplary results were correlated with an analytical model thatassesses the fluid flow characteristics through a single microneedle,which can subsequently be used to assist in the design and developmentof other embodiments of the disclosed multi-section microneedle arraystechnology, e.g., which can be applied for practical body-worn devicesthat can deliver on demand different therapeutic agents.

By employing the disclosed biosensor-actuator technology, a unique drugtherapy can be released at each microneedle constituent of the array,thereby enabling custom-tailored dosages of medications. The describedbiosensor-actuator technology includes an active solid-state device thatrequires no moving parts or integrated microelectromechanical systems(MEMS). Thus, this simplifies low-profile device design and eliminatesthe need for sophisticated microfluidics-based components, which cancomplicate system architecture and increase both size and cost.Additionally, for example, the described microneedle multi-drug deliverytechnology can be implemented with implantable devices, and thus iswell-positioned to serve as the core component in an autonomous‘wearable nanopharmacy’ in connection to multiplexed microneedle sensorarrays.

Exemplary materials and methods to implement the disclosed embodiment ofthe technology are presented. The following chemicals and reagents wereused in the described implementations, which included sodiumdodecylbenzenesulfonate (NaDBS), methylene green (MG), chresol red (CR),potassium phosphate monobasic (KH₂PO₄), and potassium phosphate dibasic(K₂HPO₄), e.g., which were obtained from Sigma Aldrich (St. Louis, Mo.)and were used without further purification or modification. Pyrrole wasdistilled daily under vacuum and stored at 4° C. prior toelectropolymerization. All reagents were prepared in a 0.1 M phosphatebuffer solution (pH 7.00). Ultrapure water (18.2 MΩ·cm) was employed inall of the exemplary implementations. Polydimethylsiloxane (PDMS) wasobtained from Dow Corning Corp. (Midland, Mich.) and mixed by hand in a10:1 polymer: fixing agent ratio. The suspension was then poured into acustom mold and degassed in a vacuum desiccator. Subsequently, the PDMSsuspension was baked at 110° C. for 15 min. The resultant structureswere exposed to UVO ozone (Jetline Co., Irvine, Calif.) at a gas flowrate of 3 sccm for 5 minutes. 25 mm-diameter black polycarbonate (PC)track etch membrane filters were procured from SPI Supplies (WestChester, Pa.); these filters possessed a pore diameter of 600 nm.

The instrumentation used in the described implementations included thefollowing, which was utilized in exemplary demonstrations andimplementations of the disclosed embodiment under exemplary conditionsdisclosed herein. A CH Instruments (Austin, Tex.) model 1232Aelectrochemical analyzer was employed for all of the electrochemicalmeasurements. An Ag/AgCl wire reference electrode and a platinum wirecounter electrode were used to establish a three-electrodeelectrochemical system. A Shimadzu (Kyoto, Japan) UV-2450 UV-VISspectrophotometer was used for all of the optical measurements. Aconsumer digital video camera/camcorder was employed to capture thestill-frame images and videos. A Philips XL30 field emission scanningelectron microscope (Amsterdam, the Netherlands) was employed toinvestigate the surface morphology of the microneedle array. The arrayswere coated with a gold film (e.g., ˜15 nm) using an Emitech (EastSussex, UK) K575X sputtering instrument prior to SEM imaging. Theresultant electron micrographs are shown in FIGS. 13A and 13B. FIG. 13Ashows an SEM image detailing the surface morphology of an exemplaryhollow microneedle array. FIG. 13B shows an enhanced view of thescanning electron micrograph of a single needle of the exemplary hollowmicroneedle array featuring a well-defined cylindrical lumen. The PCmembranes were sputtered with a gold thin film (˜75 nm) using an EmitechK 575 X sputtering instrument prior to the deposition of the NaDBS-dopedPPy conducting polymer.

The fabrication of the exemplary hollow microneedle arrays used in thedescribed implementations was performed in the following manner. Thehollow microneedle arrays were fabricated, in which, the microneedledesigns were originally prepared using Solidworks (Dassualt SystemesS.A., Velizy, France). Substrate support structures were subsequentlycreated with Magics RP 13 (Materialise NV, Leuven, Belgium). Forexample, the hollow needles were pyramidal in shape with a triangularbase. For example, the dimensions of each hollow microneedle were asfollows: an edge length of 1174±13 μm, a height of 1366±15 μm, and avertical cylindrical bore of 342±5 μm diameter on one of the faces ofthe pyramid structure. The hollow needles were arranged into 3×3 squarearrays with 2 mm periodicity. For example, substrates for themicroneedle arrays were 10 mm×10 mm in extent and possessed thicknessvalues of 500 μm.

The preparation of the exemplary electrically-actuatable nanoporousmembranes (e.g., PC/Au/PPy/DBS membranes) used in the describedimplementations was performed in the following manner. For example,gold-sputtered PC membranes (PC/Au) (e.g., pore diameter ˜600 nm,porosity ˜0.2) were attached at the periphery to a copper wire usingsilver conductive epoxy. A solution of 0.1 M NaDBS was purged withnitrogen for 40 min after which the pyrrole monomer was added to achievea final concentration of 0.25 M. Subsequently, the PC/Au membrane wasimmersed in the solution and served as the working electrode in anelectrochemical cell while 0.6 V vs. Ag/AgCl was applied for 10 min. Theapplication of this exemplary potential for the given amount of timeresulted in optimal deposition of the polypyrrole polymer on the PC/Aumembrane, thereby minimizing the leaching of the solution through themembrane under the ‘closed’ state while enabling the solution to flow atappreciable rates under the ‘open’ state. Followingelectropolymerization of polypyrrole/DBS (PPy/DBS), the PC/Au/PPy/DBSmembranes were rinsed with deionized water and stabilized by cyclingbetween −1.1 V and 0.5 V vs. Ag/AgCl for ten iterations in the buffersolution. This process enabled the membrane to swell in the reducedstate (−1.1 V) and contract in the oxidized state (0.5 V) in areversible manner. When not in use, the membranes can be stored in thebuffer solution at room temperature.

The fabrication of the exemplary drug delivery actuator contingent ofthe disclosed biosensor-actuator device used in the describedimplementations was performed in the following manner. The PC/Au/PPy/DBSmembranes were cut into slivers possessing dimensions of approximately12 mm×4 mm. These slivers were subsequently affixed to the reverse sideof the exemplary 3×3 microneedle array using adhesive epoxy such thatone sliver completely covered a column of three microneedles. The centercolumn of the array was obstructed using modeling clay, enablingformation of two individually-addressable electrically-actuatablechannels, exemplified in the component-level schematic illustrated inFIG. 12A. Electrical leads were attached using silver epoxy to each ofthe two PC/Au/PPy/DBS membranes to facilitate ohmic contact with eachactuator. The PDMS dual-channel reservoir was subsequently aligned overthe membranes and affixed using adhesive epoxy. As shown in FIG. 12B,the reservoirs were finally loaded with ˜20 μL of the model chemicalagent(s).

Initial implementations of the exemplary microneedle array actuatordevice were aimed at validating and visualizing the switching capabilityof the PC/Au/PPy/DBS membrane and the dual-channel operation. Forexample, both reservoirs in the assembled multiplexed drug deliveryactuator, e.g., reservoir 1 (R1) and reservoir 2 (R2), were initiallyloaded with 12 mM of methylene green (MG) dye and immersed in a buffersolution along with the counter and reference electrodes. Continuousagitation at a constant speed (e.g., 140 rpm) was applied with amagnetic stirring bar. The DBS-doped PPy membrane entered the reducedstate and engorged upon biasing with −1.1 V vs. Ag/AgCl, therebyobstructing the flow of the solution through the porous material.Ejection of the MG dye at either channel was not observed at thispotential (represented as being in the ‘OFF’ state), as shown in image(A) of FIG. 14. Subsequently, the R2 membrane nanoactuator wasmaintained at the reduced state (−1.1 V vs. Ag/AgCl, ‘OFF’) and themembrane at R1 was switched to the oxidized state ('ON') by applying apotential of 0.5 V vs. Ag/AgCl. This “ON” state caused the DBS-doped PPymembrane to become oxidized and contract, thereby facilitating the flowof the solution through the nanoporous membrane and subsequently throughthe microneedles. As can be observed from the image (B) of FIG. 14, theemission of MG from R1 is visible whereas R2 remained closed and did notpermit the release of the dye. Following this operation, R1 was kept atthe oxidized state (0.5 V vs. Ag/AgCl, ‘ON’) while R2 was switched tothe oxidized state (0.5 V vs. Ag/AgCl, ‘ON’), thus releasing MG fromboth reservoirs (as shown in image (C) of FIG. 14). Subsequently, R1 wasswitched to the reduced state “OFF” and R2 was kept at the oxidizedstate “ON”, as shown in image (D) of FIG. 14. This controlled andalternating release of MG from the individual reservoirs by switchingpotentials on the nanoporous membranes was illustrated in a real-timemanner. The execution of repeated ‘ON-OFF’ cycles demonstrates that thedrug delivery array maintains its ability to open and close in a cyclicfashion, e.g., even following 10 iterations or more. Furthermore, thetemporal duration (˜30 s) required to observe the release of MG at thetenth cycle was identical to that of the first cycle. In the exemplaryimplementations, the time duration for complete flow shutoff wasapproximately 35 s following the application of the “OFF” potential.Based on the above results, R1 was loaded with CR dye and R2 was loadedwith MG dye. All four “ON/OFF” permutations were applied. The controlledejection of dye from alternating microneedle array reservoirs wasdemonstrated based on the potential applied to each nanoporous membrane.

The exemplary implementations included image analysis and UV-Visspectrophotometry techniques to analyze the drug delivery capability ofthe microneedle array by experimentally quantifying the flow rate of theMG dye from a single microneedle channel. FIG. 15 illustrates therelease of MG from a single microneedle into a quiescent buffer solutionat fixed time intervals of 30 s. FIG. 15 shows exemplary time-lapsestill frame images of the release of methylene green (MG) from a singlemicroneedle at distinct time intervals of 30 s (shown in image (A)), 60s (shown in image (B)), 90 s (shown in image (C)), and 120 s (shown inimage (D)). For example, a potential of 0.5 V (vs. Ag/AgCl) was appliedto open the nanoporous membrane and release the dye during theimplementations. The exemplary flow rate of released dye was determinedto be 6.3±0.4 μL/hour (n=10) through analysis of multiple time-lapsevideo still-frames. After 30 s of applying this potential, the dye beganto emerge from the microneedle aperture. A small column of dye wasclearly observed at 60 s. A well-defined column of dye possessing aheight of approximately 0.5 cm was observed after 120 s. Afterwards, theestimated experimental flow rate of the released dye was calculated bymeasuring its column height (h) with image processing software (e.g.,ImageJ) in conjunction with the flow rate equation (Eq. 1):

$\begin{matrix}{Q = \frac{\pi\; d^{2}h}{4\left( {t - t_{0}} \right)}} & (1)\end{matrix}$where d is the microneedle channel diameter and h is the column heightassociated with a particular point at time t.

UV-Vis spectrophotometry was employed to quantify the amount of releaseddye and subsequently assess the microneedle flow rate as well as therepeatability of the release. FIG. 16 shows an exemplary UV-Vis spectrumdata plot illustrating the absorbance for the release of methylene green(MG) dye from a single microneedle at a 2 minutes release interval overa 20 minute period. The upper (top left) inset data plot displays theUV-Vis spectra, and the lower (bottom right) inset data plot displaysabsorbance of 10 distinct experimental implementations over a constanttime release. The lower (bottom right) inset data plot in FIG. 16substantiates the reproducibility of the MG release from thedrug-delivery nanoactuator over the same release time. The maximumdeviation among these ten repetitions was 5.5% from the originalabsorbance, which was measured at the maximum wavelength. Linearregression analysis was performed on the absorbance vs. time plot,yielding a slope of 3.5 mOD min⁻¹ with a high coefficient ofdetermination (R²=0.993) and low relative standard deviation (RSD=2.74%,n=3); this result indicated a constant release of dye from themicroneedle. From these implementations, the fluid flow rate wascalculated to be 5.5±0.2 μL/hour, which is in good agreement with theimage analysis data collected from the time-lapse video still-frames.The fabricated membranes exhibited excellent reproducibility, e.g.,calculated flow rates deviated by less than 10% under identicalelectropolymerization conditions.

The understanding of the fluid flow characteristics of the microneedlearray is important for delivering the precise amount of drug tosubcutaneous tissue during transdermal drug delivery. In some examples,to augment this understanding and to analytically estimate the drugdelivery capability, the fluid flow characteristics of a singlemicroneedle can be modeled via the Modified Bernoulli Equation (Eq. 2):

$\begin{matrix}{{\frac{P_{1}}{\rho g} + \frac{V_{1}^{2}}{2\; g} + z_{1}} = {\left( {\frac{P_{2}}{\rho g} + \frac{V_{2}^{2}}{2\; g} + z_{2}} \right) + {\sum\;{f\frac{{LV}_{2}^{2}}{D\; 2g}}} + {\sum\;\frac{{EV}_{2}^{2}}{2\; g}}}} & (2)\end{matrix}$where P₁ and P₂ are the atmospheric and microneedle outlet pressure, V₁and V₂ are the average fluid velocities, z₁ and z₂ are the heights atthe top of the reservoir and microneedle outlet respectively, f is thefriction factor, ρ is the fluid density, L is the channel or porelength, and D is the hydraulic diameter. FIG. 17 shows an exemplaryschematic of a single microneedle 1700 during drug delivery. Theschematic shows the following exemplary microneedle componentsincluding: a reservoir 1701 (e.g., which can store a chemical agent,such as a drug), a lumen structure 1702 (e.g., which can be a duct orcavity of a tubular structure, sized to a 342 μm diameter), a hollowmicroneedle structure 1703, an electrically-actuatable nanoporousmembranes 1704 (e.g., Au/PPY/DBS nanoporous membrane), a PC membrane1705, and the released chemical agent exiting the lumen 1706.

The second term in Eq. 2 refers to the friction losses through theactuating nanopores, polycarbonate membrane, and microneedle channel, asshown in expanded form (Eq. 3):

$\begin{matrix}{{\sum\;{f\frac{L}{D}}} = {{f_{pores}\frac{L_{pores}\;}{D}\frac{T_{pores}}{T_{pores}}} + {f_{membranes}\frac{L_{membrane}}{D}\frac{T_{membrand}}{T_{membrane}}} + {f_{microneedle}\frac{L_{microneedle}}{D}}}} & (3)\end{matrix}$where τ and c represent the tortuosity and porosity of the nanopores andpolycarbonate membranes, respectively and D is the diameter of a singlemicrochannel (e.g., 342 μm). The porosity of the PC membrane 1705 can beconfigured to 0.1 and the porosity of the actuating nanopores can beconfigured to 0.4, e.g., due to the pore narrowing created by theAu/PPy/DBS functionalization. For example, an approximate tortuosityvalue of 1.5 was assigned to the PC membrane and the actuating nanoporesto take into account the increased channel curvature created by thenanopores. The respective friction factors were calculated according toStokes flow theory for water flow in microchannels, where the product ofthe friction factor and Reynolds number (fRe=64) utilized for macroscalelaminar flow in circular channels is employed. The friction factors foreach flow section can be obtained by the Reynolds numbers obtained forfluid flow in each of the three flow sections of the microneedlechannel, as shown in Table 1. Table 1 shows dimensions and flowcharacteristics of a single microneedle channel.

TABLE 1 Length Total Cross Flow Section (μm) Sectional Area ReMicroneedle Channel 1366 A_(c) 6 × 10⁻³ Polycarbonate Membrane 7 (0.2)A_(c) 5 × 10⁻⁵ Nanoporous Membrane 0.75 (0.4) A_(c) 9 × 10⁻⁵

The values presented in Table 1 are calculated according to thefollowing exemplary considerations. For example, A_(c) is the crosssectional area of the microneedle [π(D_(microneedle))⁴/2]. The Reynoldsnumber (Re=ρVD/μ) was calculated using the density (ρ=1000 kg/m³) andviscosity (μ=1.000 N s/m²) of water at room temperature. The velocityused to estimate the Reynolds number for fluid exiting the microneedlechannel was determined a priori by averaging the experimental velocitiesobtained by image analysis and UV-Vis spectrophotometry. Furthermore,the a priori velocities for the polycarbonate membrane and thenanoporous membrane were obtained by utilizing Conservation of Mass forincompressible fluids (V₁A₁=V₂A₂) in conjunction with the averageexperimental velocity to calculate the corresponding Reynolds numbers.

The last term (ΣK) in Equation (2) represents the sum of minor lossesdue to the inlet, exit, and hydrodynamic development length, which isshown in expanded form below:ΣK−K _(inlet) +K _(outlet)  (4)where K_(inlet) and K_(outlet) are loss coefficient factors for a squareedge inlet (0.5) and for an exit (1) typically associated with hollowmicroneedles.

An expression (Eq. 5) for the theoretical flow rate of the fluid exitingthe microneedle channel can be formed by assuming quiescent flow at thetop of the reservoirs (V₁=0), and negligible pressure gradientsthroughout the flow network (ΔP=P₁−P₂=0):

$\begin{matrix}{Q_{2} = {A_{o}\sqrt{\frac{3{g\left( {z_{1} - z_{2}} \right)}}{{\Sigma\; f\frac{L}{D}} + {\Sigma\; K}}}}} & (5)\end{matrix}$

The theoretical flow rate calculated by Eq. 4 and the experimental flowrates obtained through image analysis and UV-Vis spectrophotometry werein good agreement, e.g., validating the veracity of the microneedlefluid flow model presented herein, as shown in Table 2. Table 2 exhibitsa comparison of calculated theoretical and experimental microneedle flowrates.

TABLE 2 Flow Rate (Q₂) (μl/hr) Theoretical Model 6.4 Image Analysis 6.3± 0.4 UV-Vis Spectrophotometry 5.5 ± 0.2

The ability to transdermally release multiple drugs may be important forthe autonomous treatment of metabolic syndromes (e.g., a combination ofhypertriglyceridemia, hypertension, and insulin dependent diabetesmellitus), human immunodeficiency virus, and other chronic medicalconditions. The disclosed embodiment presents a self-containedmultiplexed drug delivery system that utilizes arrays of microneedlescoupled with conducting polymer nanoactuators for the controlled releaseof fluidic agents. The ability of the exemplary PPy/DBS conductingpolymer to undergo volumetric changes with applied electrical potentialspermits the release of fluid in a controlled and switchable fashion,e.g., without the need for moving parts or integratedmicroelectromechanical systems. These nanopore-actuated microneedlearrays are well suited for integration into wearable drug deliverydevices, in which cost and power constraints must be minimized.

For example, a method to sense an analyte and deliver a therapeuticagent is described, e.g., which can be implemented using the describeddevices and systems of the disclosed embodiment. The exemplary methodcan include a process to detect a signal produced by an analyte at aninterface with a functionalized probe configured to electrochemicallyinteract with the analyte within a biological fluid, in which the signalis transduced to an electrical signal by the functionalized probe. Forexample, the functionalized probe can be one of an array of multiplefunctionalized probes, and the functionalized probe can be chemicallyfunctionalized to interact with one or more target analytes in thefluid. For example, the biological fluid can include at least one oftransdermal fluid, intraocular fluid, vitreous humor, cerebrospinalfluid, extracellular fluid, interstitial fluid, plasma, serum, lacrimalfluid, saliva, perspiration, mucus, or blood, among other biologicalfluids in a living organism. The exemplary method can include a processto process (e.g., implementing signal processing techniques) theelectrical signal to determine a parameter of the analyte (e.g., such asthe concentration of the analyte). The exemplary method can include aprocess to, e.g., based on the determined parameter, apply an electricalstimulus to a valve (e.g., in which the valve can be a porous polymerfilm having pores of a reversibly tunable porosity, as describedherein), in which the valve attached to a container containing atherapeutic agent. The exemplary method includes the electrical stimulusaltering the permeability of the pores, e.g., from a closed state to anopen state, thereby releasing the therapeutic agent into the biologicalfluid. For example, the therapeutic agent can include, but is notlimited to, a drug, vaccine, hormone, vitamin, anti-oxidant, orpharmacological agent.

The disclosed multiple-drug delivery microsystem can be integrated withthe described microneedle sensor array, e.g., coupling multiplexedanalyte detection with the corresponding therapeutic intervention. Thiscan enable a closed-loop sensing/drug delivery microneedle paradigm thatis well-positioned to precisely deliver multiple therapeutic agents inan on-demand basis. This type of autonomous “Sense-Act-Treat” system,devices, and methods can provide an avenue for responding to biomarkerfluctuations with a targeted therapy, as well as provide self-regulatingdrug delivery that can adjust patient dosage based on the severity ofthe injury or the disease process. The development of such responsivemultiplexed drug-delivering systems can be implemented to transformoutpatient, home-based civilian medical treatments as well as militarymedical care.

In another embodiment of the disclosed technology, a minimally-invasivemulti-component microneedle device with carbon paste electrodes within ahollow microneedle array for electrochemical monitoring and biosensing,which can be fabricated using a digital micromirror device-basedstereolithography techniques, is described. This embodiment can comprisethe same embodiment(s) like those previously described, and cantherefore implement the entirety of functionalities of the individualembodiments on a single embodiment. In this embodiment, a rapidprototyping method to fabricate exemplary microneedle biosensor-actuatordevices is described that uses a dynamic mask (e.g., such as a DigitalMicromirror Device (DMD)). In some examples, the exemplary method canemploy the dynamic mask for selective polymerization of a photosensitiveacrylate-based polymer resin into an exemplary microneedlesensor-actuator device. Exemplary implementations were performed thatdemonstrated that the hollow microneedles remained intact afterpuncturing the outermost layer of skin in a living organism. Forexample, in these exemplary implementations, the carbon fibers underwentchemical modification in order to enable detection of hydrogen peroxideand ascorbic acid; electrochemical measurements were demonstrated usingintegrated electrode-hollow microneedle devices. The disclosedtechnology includes an approach for implementing real-time, minimallyinvasive point-of-care sensing using an exemplary device capable ofobtaining biological samples (e.g., interstitial fluid) through the skinwhile protecting the sensing transducer from biofouling elements. Suchdevices can be used as in vivo sensors to provide real-time detection ofphysiological processes, such as monitoring of a neurotransmitters,medically-relevant molecules, cancer biomarkers, and pathogenicmicroorganisms.

Exemplary materials to fabricate and implement the disclosed embodimentof the technology are presented. For example, a variety of materials canbe used for microneedle fabrication, including silicon, glass, metal(e.g., stainless steel and nickel), and resorbable polymers (e.g.,polyglycolic acid and polylactic acid). In one example, anacrylate-based polymer, e.g., e-Shell 200, was utilized for microneedlefabrication. The material is a Class-IIa biocompatible, water-resistantmaterial; it has been used in thin-walled hearing aid shells, solidmicroneedle arrays, as well as nonmedical applications. e-Shell 200contains 0.5-1.5% wt phenylbis(2,4,6 trimethylbenzoyl)-phosphine oxidephotoinitiator, 15-30% wt propylated 2) neopentyl glycoldiacrylate, and60-80% wt urethane dimethacrylate. Energy-dispersive X-ray spectroscopyindicated that e-Shell 200 contains carbon, oxygen, and titanium; theseelements are known to possess excellent biocompatibility. e-Shell 200exhibits a water absorption value of 0.12% (D570-98 test method) and aglass transition temperature of 109° C. (E1545-00 test method). Itexhibits tensile strength of 57.8 MPa (D638M test method), flexuralstrength of 103 MPa (D790M test method), and elongation at yield of 3.2%(D638M test method).

For example, a method of fabrication included use of a DigitalMicromirror Device-stereolithography instrument to fabricate hollowmicroneedles and the integration of carbon fiber electrodes within thebores of these hollow microneedles. The carbon fibers can be chemicallymodified to enable detection of two medically significant molecules,hydrogen peroxide and ascorbic acid. Electrochemical characterizationwas performed on the chemically modified electrode-hollow microneedledevices. For example, hydrogen peroxide (H₂O₂) is a reactive oxygenspecies that is monitored in many common enzyme-based electrochemicalsensors. For example, H₂O₂ and gluconolacone are produced in reactionsbetween glucose and glucose oxidase; monitoring of released hydrogenperoxide is used for quantification of glucose. For example, monitoringof released hydrogen peroxide may also be used for quantification ofglutamate in brain dialysate; hydrogen peroxide is produced in reactionsbetween glutamate and glutamate oxidase. Glutamate is an excitatoryneurotransmitter, e.g., which has been linked with aggressive activity.Ascorbic acid can be an indicator of oxidative stress that isexperienced by cells.

The following processes were implemented to demonstrate the disclosedembodiment of a microneedle array device and fabrication methods thereofExemplary implementations included proliferation of human dermalfibroblasts and neonatal human epidermal keratinocytes on e-Shell 200surfaces, which was evaluated using the MTT(3-(4,5-dimethylthiazol-2-yl)2,5-diphenyl tetrazolium bromide) assay,e.g., which involves reduction of a yellow tetrazolium salt to a purpleformazan dye by mitochondrial succinic dehydrogenase. In the describedexemplary implementations, e-Shell 200 wafers (diameter=15 mm,thickness=2 mm) were compared against glass cover slips (diameter=15mm). The cover slips and e-Shell 200 wafers were rinsed and sterilizedin two 30 minute rinses of 70% ethanol; the materials were subsequentlyrinsed in sterile deionized water. The e-Shell 200 wafers were placed insterile Petri dishes in a laminar flow cabinet and sterilized withultraviolet B light, e.g., both surfaces were exposed to ultraviolet Blight. The materials were rotated 90 degrees after a minimum of twohours light exposure. Polymers were transferred to sterile 24-wellculture plates, rinsed twice in sterile Hank's Balanced Salt solution,and once in the appropriate cell culture medium. The e-Shell 200 waferswere placed in 2 mL of the appropriate cell culture medium and held inthe incubator until seeded.

Cryopreserved neonatal human epidermal keratinocytes (HEK) and humandermal fibroblasts (HDF) were obtained, and fibroblast growth media(FGM-2) and keratinocyte growth media (KGM-2) were also obtained (e.g.,Lonza, Walkersville, Md.). The human dermal fibroblasts and neonatalhuman epidermal keratinocytes were propagated in 75 cm² flasks, e.g.,grown to 75% confluency, and subsequently harvested. The cells wereseeded (e.g., concentration=40,000 cells per well) in a 24-well plate one-Shell wafers 200 (n=4), glass cover slips (n=4), and polystyrene wellplates (n=4). Material rinsing and all media changes were performed bymoving the test materials from one solution to the other using aforceps. The materials were placed in fresh medium after 48 hours; thistime point corresponded with 80% confluency for human dermal fibroblastsand neonatal human epidermal keratinocytes. MTT viability was assessed24 hours later. The materials with cells were rinsed using Hank'sBalanced Salt solution; desorption using isopropyl alcohol and agitationwere subsequently performed. Isopropyl alcohol (e.g., quantity=100 μL)was transferred to a new 96-well plate. Absorbance was determined (e.g.,λ=550 nm) with a Multiskan RC plate reader (Labsystems Inc., Franklin,Mass.). The mean values for percent viability were calculated.Significant differences (p<0.05) were determined using the PROC GLMProcedure (SAS 9.1 for Windows) (SAS Institute, Cary, N.C.). Whensignificant differences were found, then multiple comparisons wereperformed using Tukey's Studentized Range HSD (Honestly SignificantDifference) test at p<0.05 level of significance.

Arrays of hollow microneedles were fabricated from three-dimensionaldrawings that were created using Solidworks (Dassualt Systemes S.A.,Velizy, France). Support structures were fabricated fromthree-dimensional drawings that were created using Magics RP 13(Materialise NV, Leuven, Belgium). In the tetrahedron-shaped microneedledesign, two faces of the microneedle exhibit a vertical orientation withrespect to the substrate. The microneedle input dimensions included atriangular base with 1.2 mm sides, a height of 1.5 mm, and a verticalcylindrical channel (diameter=400 μm). The needles were arranged into a2×2 square array with 2 mm inter-microneedle spacing. The substrateinput dimensions included lateral dimensions of 1 cm×1 cm and athickness of 500 μm. Rapid prototyping of the microneedle array wasperformed using a Perfactory III SXGA+instrument (EnvisionTEC GmbH,Gladbeck, Germany). A 150-W halogen bulb was used as the light sourcefor polymerization of liquid e-Shell 200 resin. Selective polymerizationof the resin in the X-Y plane was achieved using Digital MicromirrorDevice (DMD) optics (Texas instruments, Dallas, Tex.), specifically aDMD SXGA+guidance chip with 1280×1024-pixel resolution. This instrumentcontains a build envelope of 90 mm×67.5 mm. After fabrication, themicroneedle array was washed in isopropanol in order to removeunpolymerized material. Post-building curing was accomplished using anOtoflash Post Curing System instrument (EnvisionTEC GmbH, Gladbeck,Germany), which contains two photo-flash lamps and provides lightexposure over a wavelength range of 300-700 nm.

A Hitachi S-3200 (Hitachi, Tokyo, Japan) variable pressure scanningelectron microscope with a Robinson backscattered electron detector wasused for imaging the exemplary microneedle arrays. The exemplarymicroneedle arrays were coated with 60% gold-40% palladium using aTechnics Hummer II instrument (Anatech, Battle Creek, Mich.) prior toimaging. Skin penetration testing was performed with full-thicknesscadaveric porcine skin since human skin and porcine skin exhibit similarstructures. Trypan blue (Mediatech, Inc., Manassas, Va.), atoluidine-based dye, was used to assess the transdermal drug deliveryfunctionality of the hollow microneedle arrays. Cadaveric full-thicknessweanling Yorkshire/Landrace skin was stored at 3° C. until testing wasperformed. Hollow microneedle arrays were inserted into full-thicknessporcine skin. After removal of the arrays, Trypan blue was applied tothe insertion site; the site was subsequently washed with isopropanolswabs. The Trypan blue-treated skin was subsequently imaged usingoptical microscopy. Images of a microneedle device before insertion intoporcine skin and after insertion into porcine skin were obtained usingoptical microscopy.

FIGS. 18A-18D show illustrative schematics showing processing steps forassembly of an exemplary microneedle array device of the disclosedembodiment. In this embodiment, the disclosed microneedle array deviceincludes two layers. An upper layer includes the microneedle array, andthe lower layer provides support for the carbon fibers and facilitatesalignment of the carbon fibers to the microneedle array. For example,the support component can be fabricated from a 1.6 mm thickpoly-(methylmethacrylate) (PMMA) piece. An array of holes can be laserdrilled through the PMMA piece using a Model PLS instrument, including a6.75 60-watt CO₂ laser and a computer-controlled XY stage (UniversalLaser Systems, Scottsdale, Ariz.), as shown in FIG. 18A. The holes canbe placed in a square pattern with 2 mm spacing. For example, using aModel HPDFO high power density focusing optics lens (Universal LaserSystems, Scottsdale, Ariz.), the diameter of the exemplary hole at theexit surface was measured at ˜45 μm. For example, to control the carbonfiber length beyond the support surface, the support component wasplaced on top (exit-side down) of a well with a depth of 100 μm. Thecarbon fiber can be inserted into each of the holes (entrance-side) andallowed to rest at the bottom of the well, as shown in FIG. 18B. Thefibers can be secured in place with acrylic adhesive on the entranceside after a desired well depth has been achieved, as shown in FIG. 18C.FIGS. 19A and 19B show optical images of an array of carbon fiberelectrodes and a single carbon fiber electrode in focus, respectively.The support layer and microneedle layer can be brought together in sucha manner that the carbon fibers are positioned within the hollow shaftsof the microneedles. The layers can be subsequently adhered to eachother. For example, metallic epoxy can be applied to the back of thefibers in order to create the connection for the working electrode, asshown in FIG. 18D.

The exemplary implementations included 7 μm carbon fibers (Alfa Aesar,St. Louis, Mo.) that were activated in a KOH solution (concentration=0.1M) at a pH of 13 and at a potential of 1.3 V for five minutes. In situdiazotation of 2-amino-4-nitrophenol was performed by mixing a solutionof 8 mM sodium nitrite and 6 mM 2-amino-4-nitrophenol on ice for 5minutes to create the corresponding diazonium salt. After five minutes,the activated carbon fibers were inserted. Two cyclic voltammetry (CV)scans were run from 0.4 V to −0.8 V at 0.1 V/s to enable electrochemicalgrafting of the 2-nitrophenol and subsequent reduction to theaminophenol. The carbon fibers were modified with palladium to enabledetection of hydrogen peroxide. Activated carbon fiber bundles wereplaced in a solution of 1 mM palladium (II) chloride; Pd was depositedby applying a potential of −0.8 V for 120 s. The electrochemicalmeasurements were obtained using a PGSTAT12 Autolab electrochemicalinstrument (EcoChemie, Utrecht, the Netherlands). Data was acquiredversus an Ag/AgCl reference and a Pt counter electrode.

The Digital Micromirror Device-based stereolithography instrument wasemployed for the fabrication of approximately 200 arrays over athree-hour period. FIGS. 20A and 20B show SEM images of an exemplaryhollow microneedle array and an exemplary single hollow microneedle ofthe disclosed embodiment prior to incorporation of carbon fiberelectrodes, respectively. Measurements obtained from the SEM imagesshowed that the exemplary microneedles exhibited heights of ˜1030 μm,triangular bases with side lengths of ˜1120 μm, and bore diameters of˜375 μm. Good microneedle-to-microneedle uniformity was noted among themicroneedles in the microneedle array. In this exemplary implementation,it is noted that differences between input and measured dimensions maybe attributed to translation of the computer-aided design drawing by thesoftware. For example, Digital Micromirror Device-basedstereolithography and other rapid prototyping techniques involvetessellation, conversion of the surface of the computer-aided designdrawing into a series of polygons. This polygon series is converted intoa series of cross-sectional layers, which is subsequently used forlayer-by-layer fabrication of the microneedle device. It is not possibleto predict how the computer-aided design drawing is manipulated by thesoftware.

For example, microneedles undergo bending forces, compressive forces,shear forces, and skin resistance during skin insertion; the pressurenecessary for human skin penetration can exceed 3.0×10⁶ Pa. The skinpenetration properties of the microneedle devices were evaluated usingcadaveric porcine skin, which has been previously used as a model forassessing microneedle functionality. FIG. 21A shows an image of porcineskin after application of the microneedle array, removal of themicroneedle array, and application of Trypan blue. The Trypan blue spotsindicate penetration through the stratum corneum layer (outermost layer)of the epidermis by the microneedle array and localization of TrypanBlue within microneedle-generated pores. FIG. 21B and FIG. 21C showoptical micrographs of hollow microneedles before insertion into porcineskin and after insertion into porcine skin, respectively. Theseexemplary images indicate that the microneedles remain intact after skininsertion.

For example, the positioning of the exemplary carbon fiber electrodeswithin the microneedle device can facilitate interactions with thebiological sample and minimize carbon fiber exposure to stressesassociated with microneedle insertion into skin and movement at themicroneedle device-skin interface. To facilitate interactions betweenthe biological sample and the carbon fiber electrodes, the carbon fiberelectrodes can be positioned at the centers of the microneedle bores. Inaddition, dead space between the carbon fiber electrodes and themicroneedle sidewalls may allow for infiltration of the biologicalsample. FIG. 22A shows an SEM image of a hollow microneedle array, andFIG. 22B shows an SEM image of a single hollow microneedle afterincorporation of carbon fiber electrodes. The exemplary SEM image ofFIG. 22B reveals that the carbon fiber electrodes do not extend beyondthe tip of the microneedle bore. Placement of carbon fibers within themicroneedle bores included precise alignment of the microneedle boresand the carbon fibers, e.g., the positions of the laser-ablated holes inthe lower layer of the microneedle device were coordinated with thepositions of the microneedle bores in the upper layer of the microneedledevice.

The exemplary implementation included the evaluation of theelectrochemical response of the exemplary carbon fibers within theelectrode-hollow microneedle device towards 5 mM Fe(CN)₆ ^(3/4)/1 M KCl.FIG. 23 shows a data plot of a cyclic voltammetric scan of 5 mMferricyanide in 1 M KCl versus Ag/AgCl and Pt reference counterelectrodes, respectively, e.g., at a scan rate of 100 mV/s. For example,well defined oxidation/reduction waves were observed, indicatinginteraction between the carbon fiber electrodes and the test solution asa result of permeation of the microneedle bore by the test solution. Theaverage formal potential (E^(0′)) for Fe(CN)₆ ^(3-/4-) was measured at220 mV vs. Ag/AgCl reference and platinum counter electrodes,respectively. The average peak separation was ΔE_(p)=125 mV. Theseexemplary results indicate that the carbon fibers within theelectrode-hollow microneedle device were capable of providingelectrochemical measurements.

The exemplary implementation included the evaluation ofpalladium-catalyzed oxidation of hydrogen peroxide on the carbon fiberswithin the exemplary electrode-hollow microneedle devices. For example,palladium was deposited onto the carbon fibers by applying a potentialof −0.8V for 120 sec in 1 mM Pd/0.5M HCl prior to insertion into themicroneedle device. FIG. 24 shows a data plot of cyclic voltammetricscans of 0, 50, 100, 300, and 500 μM hydrogen peroxide, e.g., as shownrepresented by pink, black, green, blue, and red curves, respectively,versus Ag/AgCl and Pt reference counter electrodes, respectively, at ascan rate of 100 mV/s. This exemplary data shown in FIG. 24 shows anincrease in reductive currents after additions of 0, 50, 100, 300, and500 μM hydrogen peroxide, exhibiting a linear range of 100-500 μM and adetection limit of ˜15 μM (based on the response of 50 μM hydrogenperoxide; S/N=3).

For example, the carbon fibers were modified with aminophenol (o-AP)groups following in-situ diazotination and electrografting of thecorresponding diazonium salt. Modification of the carbon fiber electrodewith o-AP can result in electrocatalytic oxidation of ascorbic acid andselective oxidation of ascorbic acid in the presence of commoninterferents, e.g., such as uric and citric acid. Uric acid is awell-known interferent in electrochemical analysis of ascorbic acid,e.g., which can be attributed to the fact that uric acid and ascorbicacid possess similar oxidation potential values. Linear sweepvoltammograms of 100 mM phosphate buffer (blank solution) and 1 mMascorbic acid in 100 mM phosphate buffer (pH=7) versus Ag/AgCl and Ptreference counter electrodes, respectively, at a scan rate of 100 mV/sare shown in FIG. 25. This exemplary result indicates that the carbonfibers within the electrode-hollow microneedle device are able to detectthe ascorbate analyte with the low potential oxidation of ascorbic acidat 195 mV. Electrochemical measurements by the carbon fibers within theelectrode-hollow microneedle device of the disclosed embodiment weredemonstrated. In addition, chemical modification of the exemplary carbonfibers for selective analytes was shown, and detection of hydrogenperoxide and ascorbic acid using these modified carbon fibers wasdemonstrated.

In another embodiment of the disclosed technology, a minimally-invasivemulti-component microneedle device with carbon paste electrodes (CPEs)for electrochemical monitoring and biosensing is described. Thisembodiment can comprise the same embodiment(s) like those previouslydescribed, and can therefore implement the entirety of functionalitiesof the individual embodiments on a single embodiment. The exemplarycarbon paste electrodes can exhibit a renewable nature or functionalitythat enables the packing of the exemplary hollow non-planar microneedleswith pastes that contain assorted catalysts and biocatalysts. Forexample, smoothing the surface can result inmicroelectrode-to-microelectrode uniformity. Optical and scanningelectron micrographs show the surface morphology at the microneedleapertures. Exemplary implementations of the disclosed microneedleelectrode arrays included low-potential detection of hydrogen peroxideat rhodium-dispersed carbon paste microneedles in vitro and lactatebiosensing by the inclusion of lactate oxidase in the metallized carbonpaste matrix. The exemplary implementations demonstrated highlyrepeatable sensing, e.g., for following consecutive cycles ofpacking/unpacking the carbon paste. For example, the operationalstability of the exemplary array was demonstrated, as well as theinterference-free detection of lactate in the presence ofphysiologically relevant levels of ascorbic acid, uric acid, andacetaminophen. The described microneedle design can be well-suited fordiverse biosensing applications, e.g., including subcutaneouselectrochemical monitoring of a number of physiologically-relevantanalytes.

For example, carbon paste can be characterized by a high degree ofmoldability that is essential for optimal packing and can be employed inelectroanalysis. CPEs can include the advantages of low backgroundcurrent, low cost, as well as convenient surface renewal andmodification (e.g., via the inclusion of the modifiers within thepaste). Exemplary microneedle arrays of the disclosed embodiment caninclude a nine-element arrays of pyramidal-shaped hollow microneedles,which possess a 425 μm-diameter aperture through which the modifiedcarbon paste is extruded and can act as a transducer. For example,rhodium-dispersed carbon paste, which can be used for low-potentialdetection of hydrogen peroxide, can be packed within the microneedles tominimize the contribution of co-existing electroactive interferents. Thedescribed microneedle array CPE sensor device obviates the need forintegrated microchannels and extraction of the interstitial fluid.

Exemplary materials and methods to implement the disclosed embodiment ofthe technology are presented. The following chemicals and reagents wereused in the described implementations, which included lactate oxidasefrom Pediococcus sp. (LOx, E. C. 1.13.12.4), rhodium on carbon (5% Rhweight), polyethyleneimine (PEI), mineral oil (e.g., d=0.838 g/mL),L-lactic acid, hydrogen peroxide (H₂O₂), L-ascorbic acid (AA), uric acid(UA), acetaminophen (AC), ethyl alcohol, potassium phosphate monobasic,and potassium phosphate dibasic were obtained from Sigma Aldrich (St.Louis, Mo.) and were used without further purification or modification.All experiments were performed with 0.1 M phosphate buffer (pH 7.0).Ultrapure water (e.g., 18.2 MΩ·cm) was employed in the exemplaryimplementations.

The exemplary solid and hollow microneedle arrays used in the exemplaryimplementations were developed in the following manner. The hollowmicroneedle arrays were fabricated with the aid of Solidworks (DassualtSystemes S.A., Velizy, France) computer models. Substrate structureswere designed with Magics RP 13 (Materialise NV, Leuven, Belgium). Forexample, the needles were pyramidal in shape with a triangular base. Forexample, the dimensions of each microneedle were as follows: an edgelength of 1250 μm, a height of 1500 μm, and a vertical cylindrical boreof 425 μm in diameter on one of the faces of the pyramid structure. Theexemplary needles were arranged into 3×3 square arrays with 2 mmperiodicity. Substrates for the microneedle arrays were 10 mm×10 mm inextent and possessed a thickness of 500 μm. The three-dimensionalcomputer models were transferred to a Perfactory® SXGA Standard UV rapidprototyping system (EnvisionTEC GmbH, Gladbeck, Germany) for production.This system uses these computer models to precisely guide light from a150 W halogen bulb over a photocurable material, resulting in theselective polymerization of the exposed material. Eshell 200acrylate-based polymer (EnvisionTEC GmbH, Gladbeck, Germany) wasutilized as the constituent material to fabricate the microneedle arrayssince the resin selectively polymerizes under visible light and exhibitsa Young's modulus of elasticity of 3050±90 MPa. The polymer also offersClass-IIa biocompatibility per ISO 10993. A 550 mW output power beam(e.g., step size=50 μm) with a zero-degree tilt was employed for thepolymerization of the resin. Following fabrication, the arrays wererinsed with isopropanol for removal of the unpolymerized material andsubsequently placed in an Otoflash post curing system (EnvisionTEC GmbH,Gladbeck, Germany) for post-build curing.

The exemplary enzyme-functionalized rhodium-dispersed carbon pastemicroelectrode array was prepared in the following manner. For example,100 mg of Rh-on-carbon and 10 mg of LOx were thoroughly homogenized via10 alternating 5-min cycles of vortexing and ultrasonication. Themixture was then vortexed for an additional 1 hr. Following thehomogenization process, 125 mg of the mineral oil pasting liquid and 15mg of the PEI enzyme stabilizer were added to the solid mixture.Homogenization of the resulting paste mixture was accomplished bygrinding the mixture with a mortar and pestle for an additional 1 hr.

For example, a 3 mL syringe (BD Biosciences, Franklin Lakes, N.J.) canbe utilized as the support to extrude the metallized carbon pastethrough the microneedle array. The nozzle portion of the syringe wasremoved to facilitate the attachment of the microneedle array, which wasaffixed (e.g., using adhesive epoxy) to this cleaved end for durability.A copper wire was subsequently inserted into the back end of the syringebarrel in order to create an electrical contact to the microneedletransducer. Following this exemplary procedure, the carbon paste mixturewas loaded into the syringe from the back end and then extruded with aplunger until the paste began to expel through the microneedlemicroholes. Excess paste was removed from the openings; the surface waslater smoothed using wax paper. In order to investigate therepeatability of the response after repacking the microneedles with newpaste, the array was carefully removed from the syringe and subsequentlyimmersed in ethanol under ultrasonication in order to remove theextraneous carbon paste residue. A 0.15 mm diameter iridium wire wasused to facilitate removal of the paste from the microhole. Theaforementioned assembly and packing protocols were then followed inorder to generate a new electrode from the cleaned microneedle array.

The instrumentation used in the described implementations included thefollowing, which was utilized in exemplary demonstrations andimplementations of the disclosed embodiment under exemplary conditionsdisclosed herein. A CH Instruments (Austin, Tex.) model 1232Aelectrochemical analyzer was employed for electrochemical measurements.An external Ag/AgCl reference electrode (CH Instruments CHI111) and a0.5 mm diameter platinum wire counter electrode were used to establish athree-electrode electrochemical system. The electrochemical experimentswere performed in a 7 mL cell at room temperature (22° C.). Voltammetricand chronoamperometric studies were used to evaluate the electrochemicalbehavior of the exemplary carbon paste microneedle array electrode. Inthese electrochemical implementations, either H₂O₂ or lactate was addedinto 5 mL of potassium phosphate buffer solution in order to obtain thedesired concentration. Chronoamperometric currents were sampled at 15 sfollowing the potential step. In order to obtain hydrodynamicvoltammograms, fixed potential amperograms were recorded in a stirredphosphate buffer solution containing the desired H₂O₂ concentration byvarying the potential between −0.20 and +0.60 V vs. Ag/AgCl (e.g., in0.05V increments). The solution was continuously stirred using amagnetic stirrer at a rate of 100 rpm. The morphology of the carbonpaste microneedle array was examined using a field emission scanningelectron microscope (Philips XL30, Amsterdam, The Netherlands). All ofthe specimens were coated with chromium prior to analysis using asputtering instrument (Energy Beam Sciences Emitech K575X, East Granby,Conn.). A deposition current of 130 mA was applied for 30 s to deposit˜15 nm of chromium on the sample surface.

Exemplary implementations of the disclosed embodiment includedcharacterization of the surface morphology of the carbon pastemicroelectrode array. For example, unmodified and modified carbon pastescan readily conform with the non-planar features of microneedle arraydevices. Initial implementations were aimed at characterizing themorphology of the carbon paste-loaded microneedle array and wereinitiated with a close examination of the microelectrode surface. FIGS.26A and 26B show optical micrographs of the unpacked and Rh-carbon pastepacked microneedle array, respectively. An optical micrograph of theexemplary unpacked microneedle array is shown in FIG. 26A. This imageshows uniform pyramidal microneedle structures (with triangular bases)possessing a height of 1500 μm as well as the cylindrical openings (425μm diameter). FIG. 26B depicts an exemplary microneedle array that hasbeen packed with carbon paste and subsequently polished. It indicatesthat the surface has been smoothly polished to obtain a highlyreproducible exposed area, thereby facilitating reliable electrochemicalsensing. As shown in the figures, microelectrode-to-microelectrodeuniformity can be observed.

Pursuant to the characterization of the surface morphology, SEM imagingof the microneedle was performed. FIGS. 27A and 27B show SEM images ofthe unpacked and Rh-carbon paste packed microneedle constituent of thearray. FIG. 27A shows an electron micrograph of a single microneedle.The structure of the microneedle can be observed, e.g., the boredcylindrical vacancy and the ribbed structure created by rastering of thelight source over the polymer resin. FIG. 27B shows the surface detailsof a single microneedle packed with the carbon paste. For example, asshown in the figure, a well-formed surface, a relatively smoothmorphology, and defined edges can be observed, e.g., reflecting theeffective filling of the cylindrical microhole. Such surface quality canbe achieved by extruding excess paste and later polishing the surface.It should be noted that the microneedle and the opening appear to beelongated due to the oblique angle at which the SEM image was acquired.

Exemplary implementations of the disclosed embodiment includedelectrochemical characterization of the carbon paste microelectrodearray towards peroxide-based amperometric sensing. For example, initialelectrochemical implementations were carried out to characterize theresponse of the carbon paste microneedle array to H₂O₂. A hydrodynamicvoltammogram (HDV) was recorded over the −0.20 to +0.60 V range in orderto deduce a suitable operating potential and to demonstrate the strongcatalytic ability of the Rh-CPE towards the redox processes of H₂O₂. Theexemplary results, shown in FIG. 28A, elucidate that the Rh-CPE offersconvenient detection of H₂O₂ over the entire range tested, with acrossover point occurring around 0.22V (vs. Ag/AgCl). FIG. 28A showsplots of hydrodynamic voltammograms of 0.1M potassium phosphate buffer(data plot (a)) and 10 mM H₂O₂ (data plot (b)) at the rhodium-dispersedcarbon paste microneedle electrode. For example, such lowering of theovervoltage enables the selection of a low operating potential of −0.15Vvs. Ag/AgCl for subsequent sensor implementations. At this exemplarypotential, a reduction current of 5.95 μA can be achieved for 10 mMH₂O₂. Contributions imparted by common electroactive interferences werenegligible.

The exemplary microneedle CPEs display a wide dynamic range for H₂O₂detection. FIG. 28B shows plots of chronoamperograms obtained using therhodium-dispersed carbon paste microneedle electrode (e.g., 0-500 μMH₂O₂ in 50 μM increments, a→k; E_(App)=−0.15 V vs. Ag/AgCl). Anexemplary calibration curve is shown in the inset of FIG. 28B. Forexample, as shown in FIG. 28, well-defined currents, proportional to theH₂O₂ concentration, were observed. The exemplary resulting calibrationcurve, based on sampling the current at 15 s following the potentialstep, displays high linearity (R²=0.999; as shown in the inset). Theresponse for 50 μM H₂O₂ (curve b) indicated a limit of detection (LOD)of ˜20 μM (S/N=3). The ability to detect H₂O₂ at low potentials is anattractive feature of the disclosed Rh-CPE microneedle array whenpositioned for use in minimally-invasive oxidase-based biosensors.

Exemplary implementations of the disclosed embodiment includedevaluations on the effect of reconstitution of the carbon paste matrixwithin the microelectrode array. For example, a key advantage of carbonpaste-based electrodes is their renewable surface, which can be readilyregenerated. Such regeneration should facilitate the re-use of themicroneedle array. Accordingly, the effect of repetitive packing of thearray upon the resulting response was evaluated. As such, fivecalibration experiments were executed for H₂O₂ over the 50 to 500 μMH₂O₂ range, which involved successively reconstituted carbon pastesurfaces. Between each experimental implementation, the electrode wasthoroughly disassembled, cleaned, reassembled, and repacked; itselectrochemical response was then characterized. FIG. 29 shows a plot ofa calibration curve obtained for H₂O₂ concentrations from 0 to 500 μM in50 μM increments (e.g., E_(App)=−0.15 V vs. Ag/AgCl, t=15 s). The effectof reconstitution of the Rh-dispersed carbon paste microneedle array isillustrated for five subsequent reconstitution operations. The results,illustrated in FIG. 29, are indicative of a highly-repeatablecalibration. For example, the response of successive packings deviatedby no more than 5.4% from the average current at each level over theexamined concentration range. Highly linear results were observed overthe concentration range (R²=0.997), along with a very low standarddeviation (e.g., σ<10 nA). These exemplary data demonstrate thatrepeated packing/unpacking of the carbon paste constituent in themicroneedle array resulted in a reproducible electrochemical response.

Exemplary implementations of the disclosed embodiment included thebiosensing of lactate at the microneedle CPE arrays. For example, anexemplary microneedle array CPE biosensor for lactate was developed.Lactate oxidase (LOx)-dispersed metallized carbon paste was preparedusing PEI for the electrostatic entrapment of the enzyme within thematrix. Chronoamperometric calibration experiments were performed usingthe LOx-Rh-carbon paste microneedle array at −0.15 V vs. Ag/AgCl forincreasing levels of lactate (e.g., 0 to 8 mM in 1 mM increments).Typical chronoamperograms are displayed in FIG. 30A, which showsexemplary plots of chronoamperograms obtained for lactate concentrationsfrom 0 to 8 mM in 1 mM increments (e.g., E_(App)=−0.15 V vs. Ag/AgCl).FIG. 30B shows an exemplary calibration curve corresponding to thechronoamperometric current at t=15 s. For example, high linearity(R²=0.990) and low deviation (e.g., σ<10 nA) were observed. A detectionlimit was estimated to be 0.42 mM lactate (S/N=3), which is well belownormal physiological levels and is therefore more than sufficient forrelevant applications. It should be noted that the exemplary linearconcentration range encompasses the entire physiological andpathological range of lactate in transdermal fluids, e.g., indicatingthe diagnostic value of the microneedle-based lactate biosensor.

Exemplary implementations of the disclosed embodiment includedevaluating the microneedle CPE arrays for interference by commonelectroactive compounds. For example, in order to ascertain that theexemplary biosensor could function as intended in the presence of commonelectroactive substances found in transdermal fluids, an interferenceinvestigation was conducted using physiological levels of thesecompounds. FIG. 31 shows plots of chronoamperograms showing the effectof physiologically-relevant electroactive interferents upon thedetection of 1 mM lactate in the presence of 60 μM ascorbic acid (AA),500 μM uric acid (UA), and 200 μM acetaminophen (AC) (e.g.,E_(App)=−0.15 V vs. Ag/AgCl). As shown in the figure, the addition ofany of these common electroactive interferents resulted in a negligibleeffect on the lactate response of the exemplary biosensor device. Forexample, a maximum current deviation of only 1.5% from the 1 mM lactatelevel was observed for the addition of AC. For example, suchinterference-free lactate detection reflects the strong and preferentialelectrocatalytic activity of the Rh-CPE towards H₂O₂, which can bedetected by the described microneedle paste biosensor, e.g., for lactatemonitoring in transdermal fluids.

Exemplary implementations of the disclosed embodiment includedevaluating the stability of the lactate response of the microneedle CPEarrays. For example, the stability of the microneedle array-basedbiosensor was examined from repetitive chronoamperograms for 2 mMlactate over a 2 hour period. In some examples, an initial shortpreconditioning step was implemented. This process involved theimmersion of the exemplary CPE microneedle array in a 0.1 M potassiumphosphate buffer (pH 7.0) and the concomitant recording of sixchronoameprograms, followed by the immersion of the array in a 2 mMlactate solution for 10 min while recording two chronoamperograms. Afterthe exemplary precondioning, the current was sampled every 10 min overthe entire 2 hour stability test period. FIG. 32 illustrates thetime-course profile of the resulting current response, e.g., with theinitial reading at t=0 min normalized to 100%. FIG. 32 shows a data plotof the stability of the electrochemical response of the microneedlearray for 2 mM lactate (E_(App)=−0.15V vs. Ag/AgCl) over a 2 hourduration. As shown in the figure, a stable current was achieved almostimmediately following the initialization of the experiment, with only aslight increase (of 9.7%) over the entire 2 hour time course. The stableresponse reflects the integrity of the exemplary CPE microneedle arraybiosensor. For example, tight packing of the CPE, which can prevent thepotential accumulation of the enzymatic product within the microneedleopenings, can influence the stable response.

For example, in the disclosed embodiment, the coupling of CPEtransducers with microneedle hosts was shown to provide low-potentialdetection of H₂O₂. For example, exemplary implementations of thedisclosed embodiment demonstrated that a reproducible amperometricresponse can be achieved following successive reconstitution of thecarbon paste matrix. For example, exemplary implementations of thedisclosed embodiment demonstrated that highly-linear lactate detectioncan be achieved over the entire physiological range, along with the highselectivity imparted by the very low cathodic detection potential. Thehigh selectivity, sensitivity, and stability of the described CPEmicroneedle array demonstrates the ability of the array to beimplemented in diverse on-body sensing applications.

While this patent document contains many specifics, these should not beconstrued as limitations on the scope of any invention or of what may beclaimed, but rather as descriptions of features that may be specific toparticular embodiments of particular inventions. Certain features thatare described in this patent document in the context of separateembodiments can also be implemented in combination in a singleembodiment. Conversely, various features that are described in thecontext of a single embodiment can also be implemented in multipleembodiments separately or in any suitable subcombination. Moreover,although features may be described above as acting in certaincombinations and even initially claimed as such, one or more featuresfrom a claimed combination can in some cases be excised from thecombination, and the claimed combination may be directed to asubcombination or variation of a subcombination.

Similarly, while operations are depicted in the drawings in a particularorder, this should not be understood as requiring that such operationsbe performed in the particular order shown or in sequential order, orthat all illustrated operations be performed, to achieve desirableresults. Moreover, the separation of various system components in theembodiments described in this patent document should not be understoodas requiring such separation in all embodiments.

Only a few implementations and examples are described and otherimplementations, enhancements and variations can be made based on whatis described and illustrated in this patent document.

What is claimed is:
 1. A method for in situ measurement of one or moretarget chemical species present in a physiological fluid, the methodcomprising: causing a microneedle sensing device to contact thephysiological fluid, wherein the microneedle sensing device includes asubstrate and an array of electrically-conducting electrodes, whereineach electrically-conducting electrode is respectively disposed in ahollow interior of a hollow microneedle of a hollow microneedle arrayformed on the substrate, one electrode per hollow microneedle, whereineach electrically-conducting electrode has a tip located inside thehollow interior of its respective hollow microneedle to form a recessmicrocavity that surrounds the electrode within the hollow interior,wherein the tip is coated to be functionalized for interacting with oneor more target chemical species present in a portion of thephysiological fluid that is trapped in its respective recessmicrocavity; applying a stimulus signal to an electrically conductivestructure disposed on the substrate that is coupled to the array ofelectrically-conducting electrodes to cause biosensing of thephysiological fluid entrapped in each recess microcavity; and reading aresultant signal arising at an interface between the physiological fluidand the coated tip of the electrically-conducting electrode in eachrecess microcavity as part of the in situ measurement, wherein aproperty of said resultant signal is indicative of a concentration ofthe one or more target chemical species present in said physiologicalfluid.
 2. The method according to claim 1, comprising: processing theresultant signal to determine whether the concentration of the one ormore target chemical species present in said physiological fluidreflects a healthy or disease state.
 3. The method according to claim 2,wherein processing the resultant signal includes comparing the resultantsignal to a threshold value.
 4. The method according to claim 2, whereinprocessing the resultant signal includes determining a pattern in theresultant signal.
 5. The method according to claim 1, wherein theproperty of said resultant signal includes magnitude of the resultantsignal.
 6. The method according to claim 1, wherein the property of saidresultant signal includes an impedimetric property.
 7. A device for insitu measurement of one or more target chemical species present in aphysiological fluid, the device comprising: a substrate; a hollowmicroneedle array of hollow microneedles formed on one side of thesubstrate, each hollow microneedle including a hollow interior; an arrayof electrically-conducting electrodes, wherein eachelectrically-conducting electrode is respectively disposed in the hollowinterior of one of the hollow microneedle, one electrode per hollowmicroneedle, wherein each electrically-conducting electrode has a tiplocated inside the hollow interior of its respective hollow microneedleto form a recess microcavity that surrounds the electrode within thehollow interior, wherein the tip is coated to be functionalized forinteracting with one or more target chemical species present in aportion of the physiological fluid that is trapped in its respectiverecess microcavity; an electrically conductive structure disposed on thesubstrate, wherein the conductive structure is configured to receive anexcitation signal and is electrically connected with each of theelectrodes to cause electrochemical sensing of a portion of thephysiological fluid that is trapped in each recess microcavity.
 8. Thedevice of claim 7, wherein the device is configured to measure aresultant signal at an interface between each electrode and thephysiological fluid; and wherein a property of said resultant signal isindicative of a concentration of the one or more target chemical speciespresent in said physiological fluid.
 9. The device according to claim 8,wherein the property of said resultant signal includes magnitude of theresultant signal.
 10. The device according to claim 8, wherein theproperty of said resultant signal includes an impedimetric property. 11.The device according to claim 7, wherein said array of hollowmicroneedles include at least one of a conal, columnar, triangular,pyramidal, rectangular, or polygonal geometric shape.
 12. The deviceaccording to claim 7, wherein each microneedle is made of a polymer. 13.The device according to claim 7, wherein said substrate is planar andcomprises one of a polymer, semiconductor, metal, glass, or film. 14.The device according to claim 7, wherein each tip is functionalized by achemical species-selective layer deposited on a surface of said tip andfacilitates quantification of the target species.
 15. The deviceaccording to claim 14, wherein said chemical species includes one of ametabolite, electrolyte, ion, pathogen, microorganism, hormone, protein,enzyme, vitamin, anti-oxidant, nucleic acid, antibodies,neurotransmitter, acid, or drug.
 16. The device according to claim 14,wherein the chemical species-selective layer comprises one of anion-selective membrane, enzyme-loaded membrane, conducting polymer,diffusion-limiting polymer, electrostatically-bound biocatalyst,covalently-bound biocatalyst, or surface-adsorbed biocatalyst.
 17. Thedevice according to claim 7, wherein said physiological fluid includesliquids originating from a living human body and lying beneathbiological tissues or membranes.
 18. The device according to claim 7,wherein the physiological fluid comprises one of interstitial fluid,intracellular fluid, extracellular fluid, blood, plasma, serum,cerebrospinal fluid, mucus, intraocular, or vitreous humor.
 19. Thedevice according to claim 7, wherein the electrically-conductingelectrodes comprise a platinum coating.
 20. The device according toclaim 7, wherein each microneedle extends from the substrate to adistance of between 20 to 1500 micrometers.